Estimation of blood flow and hemodynamic parameters from a single chest-worn sensor, and other circuits, devices and processes

ABSTRACT

An electronic monitoring device includes an electronic processor ( 520 ) having at least one signal input for body monitoring, and a memory ( 530 ) holding instructions for the electronic processor coupled to the electronic processor so that the electronic processor is operable to isolate a cardiac signal including cardiac pulses combined with other cardiac signal variations, and the electronic processor further operable to execute a filter ( 730 ) that separates a varying blood flow signal from the cardiac pulses and to output information ( 790 ) based on at least the varying blood flow signal. Other devices, sensor assemblies, electronic circuit units, and processes are also disclosed.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application is related to U.S. patent applications as follows:

This application is related to U.S. patent application: “Heart MonitorsAnd Processes With Accelerometer Motion Artifact Cancellation, And OtherElectronic Systems” Ser. No. 12/______ (TI-68518) filed Aug. 24, 2010simultaneously herewith, for which priority is claimed under 35 U.S.C.120 and all other applicable law, and which is incorporated herein byreference in its entirety.

This application is related to U.S. patent application “Motion/Activity,Heart-Rate And Respiration From A Single Chest-Worn Sensor, Circuits,Devices, Processes And Systems” Ser. No. 12/______ (TI-68552) filed Aug.24, 2010 simultaneously herewith, for which priority is claimed under 35U.S.C. 120 and all other applicable law, and which is incorporatedherein by reference in its entirety.

This application is related to provisional U.S. patent application“Motion Artifact Cancellation to Obtain Heart Sounds from a SingleChest-Worn Accelerometer” Ser. No. 61/242,688 (TI-68518PS) filed Sep.15, 2009, for which priority is claimed under 35 U.S.C. 119(e) and allother applicable law, and which is incorporated herein by reference inits entirety.

This application is related to provisional U.S. patent application“Motion/Activity, Heart-rate and Respiration From a Single Chest-wornSensor” Ser. No. 61/262,336 (TI-68552PS) filed Nov. 18, 2009, for whichpriority is claimed under 35 U.S.C. 119(e) and all other applicable law,and which is incorporated herein by reference in its entirety.

This application is related to provisional U.S. patent application“Estimation of Blood Flow and Hemodynamic Parameters from a SingleChest-worn Sensor” Ser. No. 61/262,331 (TI-68553PS) filed Nov. 18, 2009,for which priority is claimed under 35 U.S.C. 119(e) and all otherapplicable law, and which is incorporated herein by reference in itsentirety.

This application is related to provisional U.S. patent application“Heart Rate Detection In High Noise Conditions” Ser. No. 61/104,030(TI-66732PS) filed Oct. 9, 2008, for which priority is claimed under 35U.S.C. 119(e) and all other applicable law, and which is incorporatedherein by reference in its entirety.

This application is related to U.S. patent application Publication“Heart Rate Detection In High Noise Conditions” 20100094150, dated Apr.15, 2010 (TI-66732) for which priority is claimed under 35 U.S.C. 120and all other applicable law, and which is incorporated herein byreference in its entirety.

This application is related to provisional U.S. patent application“Robust Heart Rate Detection in the Presence of Pathological Conditions”Ser. No. 61/023,581, filed on Jan. 25, 2008 (TI-65798PS), for whichpriority is claimed under 35 U.S.C. 119(e) and all other applicable law,and which is incorporated herein by reference in its entirety.

This application is related to U.S. patent application Publication“Method and System for Heart Sound Identification” 20090192401, datedJul. 30, 2009 (TI-65798) for which priority is claimed under 35 U.S.C.120 and all other applicable law, and which is incorporated herein byreference in its entirety.

This application is related to U.S. patent application Publication“Method and Apparatus for Heart Rate Monitoring” Ser. No. 12/768,488filed Apr. 27, 2010 (TI-67877), which is incorporated herein byreference in its entirety.

This application is related to U.S. patent application “ParameterEstimation for Accelerometers, Processes, Circuits, Devices and Systems”Ser. No. 12/398,775 (TI-65353) filed Mar. 5, 2009, and which isincorporated herein by reference in its entirety.

This application is related to the US patent application titled“Processes for More Accurately Calibrating E-Compass for Tilt Error,Circuits, and Systems” Ser. No. 12/398,696 (TI-65997) filed Mar. 5,2009, and which is incorporated herein by reference in its entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

Not applicable.

COPYRIGHT NOTIFICATION

Portions of this patent application contain materials that are subjectto copyright protection. The copyright owner has no objection to thefacsimile reproduction by anyone of the patent document, or the patentdisclosure, as it appears in the United States Patent and TrademarkOffice, but otherwise reserves all copyright rights whatsoever.

FIELD OF TECHNOLOGY

The field of technology is in the areas of monitoring of the human body,automatic analysis and display of monitoring data locally for medicaland other purposes and telecommunication remotely for tele-medicine, andprocesses, circuits and devices for body monitoring of heart function,circulatory function, respiration, or other physiological processes.Biomedical instrumentation and signal processing are further fields.

BACKGROUND

Ambulatory measurement of cardiac activity can facilitate home healthmonitoring of older adults and of patients with a history ofcardiovascular conditions. Evaluating cardiovascular performance ofpatients in ICU (intensive care unit) and hospital settings, in mobileambulances, and at accident and trauma sites also involves or caninvolve ambulatory cardiac measurement.

Most current solutions for heart rate monitoring involve cumbersomeequipment, such as heart rate recording belts to be worn around thechest, electrocardiogram (ECG) electrodes and leads, and in most caseselectrical contact to the skin. However, such methods remain obtrusive,and are not optimal for long-term and ambulatory monitoring.

An alternative method of heart rate measurement uses heart sounds,conventionally measured with stethoscopes or phonocardiograph.

Detection and early warning of risk factors for and any incident ofheart failure is vitally important in medicine, allied medical fields,residential care-giving, exercise venues and other settings. Heartfailure can be caused by, and is at risk in case of, coronary arterydisease, hypertension, valve disorder, past myocardial infarction,muscle disorder, congenital heart conditions, etc.

Current solutions for not only heart rate monitoring but alsorespiration monitoring are believed to involve cumbersome and expensiveequipment e.g., respiration and heart rate monitoring belts to be wornaround the chest, spirometers and canulas to be worn around the mouthand nose, and electrocardiogram (ECG) electrodes and leads to be tapedon the body. Not only are these solutions obtrusive and expensive, butmay also be too restrictive to be well-suited for ambulatory monitoring.

Noise mixed with signals received by the sensors used in heartmonitoring, respiration monitoring, body motion and other monitoringapplications can adversely affect the accuracy of each type of signal.Accordingly, methods for robust detection and separation of such signalsin noisy conditions are desirable. Accuracy of heart rate detection isimportant in many commercial heart monitoring applications (e.g., heartrate monitors in exercise equipment, personal heart rate monitors, etc.)and medical heart monitoring applications (e.g., digital stethoscopes,mobile cardiac monitoring devices, etc.).

Simpler, more economical and more efficient methods and devices aredesirable in the art for obtaining, isolating, determining andmonitoring resting data and ambulatory data, such as robust, accuratedetection of heart rate, timings of heart sounds (S1 and S2) andpathological cardiac conditions, and robust detection of respiration inconnection with respiratory and pulmonary disorders, as well as data onbody motion and ambulatory data and activity data.

Conventional approaches to address the bodily motion signal separationand/or removal problem are believed to involve multi-signal adaptivealgorithms that need an additional motion signal reference recordingtypically from a secondary sensor. Also, the reference signal needs tobe reasonably well correlated to the motion picked up by the primarysensor. Such arrangements are very difficult to establish in a realsetting and can cause poor rejection of the motion signal and bodymotion artifacts. Some conventional single-channel de-noising techniquesreinforce all major signal peaks and fail to distinguish body motionsfrom heart sounds.

In addition to medical-related applications, solving the above problemscould also help monitor older adults for unexpected changes in gait, forfalls, for syncope (fainting), for accidents and trauma incidents.Fitness monitoring at home, in exercise venues, and in institutionalcare settings could also benefit.

Hemodynamic data also challenge the art to find methods and devices forobtaining, isolating, determining and monitoring more simply,economically and more efficiently. Hemodynamics as discussed hereinincludes the study of blood flow-related data directly or indirectlyrelated to blood flow, such as: heart stroke volume, cardiac output,pre-ejection period, contractility (ability of heart to contract,inotropy), and related causal or caused bodily dynamics such as exerciseand exercise recovery, and the Valsalva maneuver (such as when pushingor straining while holding one's breath, or otherwise doing the maneuverin a medical test).

Measurement of blood flow, hemodynamics and cardiovascular performanceis integral to a holistic assessment of an individual's health.Specifically, patients with past conditions of heart disease like heartfailure (potentially arising out of one or more of many causes likecoronary artery disease, heart valve or heart muscle disorders, pastmyocardial infarction, hypertension etc.) may need constant monitoringin order to improve a person's quality of life via timely andappropriate diagnostic interventions. While the physiological mechanismsunderlying these conditions are fairly well understood, the technologyto monitor these physiological vitals needs considerable improvement.

Most current solutions for the measurement of blood flow and otherhemodynamic parameters are believed to involve cumbersome and expensiveequipment e.g., Impedance Cardiography (calls for electrodes to beconnected on the skin), Doppler Echo Cardiography, Continuous BloodPressure Monitoring etc. Not only are these solutions obtrusive andexpensive, but may also be too restrictive to be well-suited forambulatory monitoring applications.

SUMMARY OF THE INVENTION

Generally, and in one form of the invention, an electronic monitoringdevice includes an electronic processor having at least one signal inputfor body monitoring, and a memory holding instructions for theelectronic processor coupled to the electronic processor so that theelectronic processor is operable to isolate a cardiac signal includingcardiac pulses combined with other cardiac signal variations, and theelectronic processor further operable to execute a filter that separatesa varying blood flow signal from the cardiac pulses and to outputinformation based on at least the varying blood flow signal.

Generally, and in another form of the invention, an accelerometer sensorassembly has a broadside portion and includes an accelerometer sensorcircuit having an axis of acceleration sensitivity parallel to the broadside and orientable on the chest to deliver an input signal representinga component of acceleration approximately parallel to a head-to-feetbody axis and including heart pulses and other variations mixedtogether, an electronic circuit responsive to the accelerometer sensorcircuit and operable to execute a filter that delivers a varying bloodflow signal from the input signal at least when that axis ofacceleration sensitivity is approximately parallel to the head-to-feetbody axis, the varying blood flow signal substantially freed of heartpulses, and an output circuit operable to send information based on thevarying blood flow signal.

Generally, and in a process form of the invention, an electronicmonitoring process includes electronically bandpassing digital signalsrepresenting living-body monitoring signals in a range selective for acardiac signal, and executing a filter that delivers a varying bloodflow signal from the cardiac signal.

Generally, and in a further form of the invention, a hemodynamicmonitoring device includes an electronic processor having at least onesignal input for body monitoring, and a memory holding instructions forthe electronic processor coupled to the electronic processor so that theelectronic processor is operable to obtain a cardiac pulse signal havinga varying amplitude, and the electronic processor further operable topost-process the cardiac pulse signal to provide a time-varying outputrepresenting an estimation for at least one hemodynamic parameter basedon the amplitude and selected from the group consisting of stroke volumeSV and cardiac output CO.

Other devices, sensor assemblies, electronic circuit units, andprocesses are also disclosed and claimed.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a partially-block, partially-pictorial, partially graphicaldepiction of an inventive structure and process for separating a heartsignal from body motion and noise using a single chest sensor.

FIG. 2 is a partially-schematic, partially-pictorial, partiallygraphical depiction of a measurement setup including both anaccelerometer sensor on the chest with a composite signal having heartsignals and motion and noise, as well as an electrocardiogram ECGcircuit with ECG electrodes affixed to the body and showing anoperational amplifier and an ECG signal.

FIG. 3 is a block diagram of an inventive structure and process forseparating a heart signal from body motion and noise using a singleaccelerometer chest sensor.

FIG. 4 is a flow diagram of the inventive structure and process forseparating a heart signal from body motion and noise using a singleaccelerometer chest sensor and remarkable smoothing filter and residuecircuit, envelope-based noise rejection, folded correlation and othersteps.

FIG. 5 is a set of four concurrent waveform traces of voltage versustime in various parts of the inventive structure and process of FIGS. 2,3 and 4 with a subject walking around a room. A time interval portion ofthe traces is magnified and shown as four time-magnified waveformsmaintaining the same voltage scale for each. Some of the traces areaccelerometer-based and one is ECG-based.

FIG. 6 is a pair of concurrent accelerometer-based waveform traces ofvoltage versus time in parts of the inventive structure and process ofFIGS. 2, 3 and 4 with a subject walking on a treadmill. A time intervalportion of the traces is magnified and shown as time-magnified waveformsmaintaining about the same voltage scale for each.

FIGS. 7A, 7B, and 7C are each a pair of concurrent noisy handgripECG-based waveform traces of voltage versus time in parts of theinventive structure and process of FIGS. 2 and 4 with a subject walkingon a treadmill. The ECG-based waveforms are one unfiltered, and oneinventively filtered to recover a heart sounds signal.

FIG. 8 is a block diagram of another inventive structure and process forobtaining cardiac information using an accelerometer chest sensor, suchas heart rate and information related to beat-to-beat changes in strokevolume and cardiac output.

FIG. 9 is a pair of concurrent accelerometer-based waveform traces ofvoltage versus time in parts of the inventive structure and process ofFIGS. 3 and 4 pertaining to envelope-based noise rejection.

FIG. 10 is a pair of concurrent accelerometer-based waveform traces ofvoltage versus time in parts of the inventive structure and process ofFIGS. 3 and 4 pertaining to folded correlation.

FIG. 11 is a plot of the difference of two heart-rate measures(inventively filtered accelerometer-based and ECG-based) versus theiraverage.

FIG. 12 is a plot having time interval between adjacent S1 cardiacpulses from inventively filtered accelerometer data on one graph axis,versus ECG R-R interval on the other graph axis.

FIG. 13 is a partially-block, partially-pictorial, partially graphicaldepiction of an inventive structure and process for separating arespiration signal from heart and body motion and other signals using asingle chest sensor.

FIG. 14 is a partially-flow, partially graphical depiction of aninventive process for FIG. 13 separating a respiration signal, a heartsignal and a body motion signal from each other using a single chestsensor.

FIG. 15 is a pair of concurrent accelerometer-based waveform traces ofvoltage versus time in parts of the inventive structure and process ofFIGS. 13-14 and shows a raw and inventively filtered accelerometer-basedsignal during rest and brisk motion. A time portion of the signalsduring rest is magnified in both voltage scale and time scale. A timeportion of the signals during subsequent motion is magnified in timescale and not voltage scale.

FIG. 16 is a set of four concurrent waveform traces of voltage versustime in various parts of the inventive structure and process of FIGS.13-14 during motion and brisk walking A time interval portion of thetraces is magnified and shown as four time-magnified waveformsmaintaining the same voltage scale for each. Some of the traces areinventively filtered accelerometer-based and one is ECG-based.

FIG. 17 is a flow diagram of a process for FIGS. 13-14 to separate arespiration signal from a heart signal and using inter-beat intervals ofthe heart signal, with both the respiration signal and the heart signalsubstantially separated from body motion and noise signals.

FIG. 18 is a set of three concurrent waveform traces of voltage versustime in various parts of the inventive structure and process of FIGS.13-14, showing raw signal and inventively-obtained residue from filteredaccelerometer-based signal, and further showing a respiration signalgenerated from ECG—R-R interval, and accelerometer heart sounds S1-S1interval of the residue signal. A time interval portion of the traces ismagnified and shown as three time-magnified waveforms maintaining thesame voltage scale for the first two, and magnifying the voltage scalefor the respiration signal.

FIG. 19 is a flow diagram of another process for FIGS. 13-14 to separatea respiration signal from a heart signal according to baseline wandermethod herein for respiration monitoring by a single inventivelyfiltered accelerometer sensor.

FIG. 20 is a set of four example waveforms of voltage versus time andshown for comparison of a respiration belt signal with respirationoutputs from each of the processes of FIGS. 17, 19 and 22.

FIG. 21 is a set of seven example waveforms of voltage versus time,including a comparison of a reference respiration signal and ECG-derivedrespiration signal with respiration outputs from each of the processesof FIGS. 17, 19 and 22.

FIG. 22 is a flow diagram of a process for FIG. 13 to inventivelyseparate a respiration signal from a heart signal by amplitudemodulation detection of peak heights of the heart signal.

FIG. 23 is a set of three concurrent waveform traces of voltage versustime in various parts of the structure and process of FIGS. 2 and 13 and22, showing ECG amplitude modulation on the R peaks, and further showingamplitude modulation on the S1 peaks from inventively filteredaccelerometer based sensing, and also showing a respiration signalobtained from a respiration belt for reference.

FIG. 24 is a block diagram of an inventive wired system structure andprocess and including inventive structures and processes from the otherFigures.

FIG. 25 is a block diagram of an inventive wireless system structure andprocess and including inventive structures and processes from the otherFigures.

FIG. 26 is a partially-block, partially-pictorial, partially graphicaldepiction of an inventive structure and process for separating a bloodflow signal from heart and other signals using sensor signals from oneor more axes of a single chest sensor.

FIG. 27 is a pair of concurrent accelerometer-based waveform traces ofvoltage versus time of sensor signals from multiples axes of a singlechest sensor in the inventive structures and processes of FIG. 26.

FIG. 28 is a voltage-versus-time graph of a pair of concurrentaccelerometer-based waveforms from Z- and Y-axes of the single chestsensor, along with inventively filtered Y-axis signal and residue in theinventive structures and processes of FIG. 26.

FIG. 29 is a voltage-versus-time graph of three concurrent waveformsincluding a pair of inventively filtered accelerometer-based waveformsfrom Z- and Y-axes of the single chest sensor in the inventivestructures and processes of FIGS. 26, 30 and 31, compared with areference ECG waveform.

FIG. 30 is a partially-block, partially-pictorial, partially graphicaldepiction of another inventive structure and process for separating ablood flow signal and hemodynamic parameters, respiration signals, heartsignals and motion signals from each other using sensor signals from oneor more axes of a single chest accelerometer sensor.

FIG. 31 is a combined flow diagram of inventive processes for separatinga heart signal from body motion and noise using Z-axis sensor input andseparating a blood flow signal using Y-axis sensor input from FIG. 30and further electronically processing the heart signal and blood flowsignal jointly to generate hemodynamic parameter signals for a displayas in FIGS. 24 and 25.

FIG. 32 is a voltage-versus-time graph of three concurrent waveformsincluding an ECG signal, a filtered heart signal from the Z-axisaccelerometer sensor, and a blood flow signal filtered from Y-axis ofthe accelerometer sensor in the inventive structures and processes ofFIGS. 26, 30 and 31, and further showing time locations P1 and F1 andhemodynamic parameters for isovolumic contraction interval IVCI,pre-ejection period PEP and flow peak amplitudes PAmp and Jamp.

FIG. 33 is a graph of voltage (arbitrary units) versus time-samples fora multitude (ensemble) of waveforms each of a respective instance ofinventively filtered blood flow signal from Y-axis of the accelerometersensor in the inventive structures and processes of FIGS. 26 and 30.(FIG. 33 is on same sheet as FIG. 37.)

FIG. 34 is a voltage-versus-time graph of four concurrent waveformsduring exercise recovery, the waveforms including reference ECG,inventively filtered blood flow signal from Y-axis, PEP, and PAmp. Atime interval portion of the traces is magnified and shown as fourtime-magnified waveforms maintaining the same voltage scale for eachexcept for modest voltage scale magnification for PEP and PAmp.

FIG. 35A is a voltage-versus-time graph of four concurrent waveformsover about a minute for a Valsalva Release phase of a Valsalva maneuver;the first waveform representing inventively-produced residue frompolynomial filtering the accelerometer Z-axis as in FIG. 31 (left side),the second waveform representing peak amplitude PAmp of that residue,the third waveform representing Stroke Volume, and the fourth waveformrepresenting Cardiac Output.

FIG. 35B is a voltage-versus-time graph of another four concurrentwaveforms over about a minute for a Valsalva Release phase of a Valsalvamaneuver; the first waveform representing the blood flow signal frominventive polynomial filtering of the accelerometer Y-axis as in FIG. 31(right side), the second waveform representing peak amplitude PAmp ofthat blood flow signal from accelerometer Y-axis, the third waveformrepresenting Stroke Volume, and the fourth waveform representing CardiacOutput.

FIG. 36A is a flow diagram of inventive process for separating a heartsignal from body motion and noise using Z-axis sensor input such as foruse in FIG. 31 or with FIG. 36B.

FIG. 36B is a flow diagram of inventive process for separating a heartsignal as well as a blood signal from each other using Y-axis sensorinput such as for use in FIG. 31 and for obtaining further hemodynamicdata and other information from the single Y-axis of the chest sensor.

FIG. 37 is a model of a standing subject, the model described by asecond-order differential equation to approximate the blood flow signalof the standing subject as a solution thereof.

FIG. 38 is a model of a subject lying prone, the model described by asecond-order differential equation having different model parametersthan in FIG. 37, to approximate the blood flow signal of the pronesubject as a solution thereof.

FIG. 39 is a block diagram of a system structure for use in and improvedaccording to inventive structures and processes from the other Figures.

FIGS. 40A and 40B are respective broadside and cross-sectional views ofan inventive accelerometer sensor and transponder chip mounted on asupport plate affixed by an adhesive tape to the chest, and for use withthe inventive structures and processes from the other Figures.

FIG. 41 is a block diagram of an inventive structure and process forvariably combining accelerometer signals from multiple axes in variousproportions to provide one or more inputs to the smoothing filtering ofFIGS. 31, 36A and 36B and various other Figures herein.

FIG. 42 is a voltage-versus-time graph of four concurrent waveformsincluding reference ECG, acceleration along a dorso-ventral axis(Z-axis), acceleration along a superior-inferior axis (Y-axis) andacceleration along a dextro-sinistral axis (X-axis), the variousacceleration signals for use in the circuit FIG. 41 and circuits ofother Figures.

Corresponding numerals in different Figures indicate corresponding partsexcept where the context indicates otherwise. A minor variation incapitalization or punctuation for the same thing does not necessarilyindicate a different thing. A suffix .i or .j refers to any of severalnumerically suffixed elements having the same prefix. A first, second,third, etc. waveform is referenced in top to bottom order for a givenFigure.

DETAILED DESCRIPTION OF EMBODIMENTS

Some structure and process embodiments provide motion artifactcancellation or motion signal separation to obtain heart sounds from asingle chest-worn accelerometer.

Miniature, high-sensitivity MEMS accelerometers are presently available.Here, such an accelerometer is incorporated into a single, chest-wornsensor for recording of signals including some related to heart sounds.(The latter signal components are also themselves sometimes called heartsounds herein. The term “heart sound” refers in an expansive way to asignal analogous to cardiac S1, S2, and/or heart murmur or other cardiacwaveform features, obtained from the processing of accelerometer data orother sensor data, and not necessarily to an audible sound.)

However, a major challenge of ambulatory monitoring is the corruption ofheart signals by body motion artifact signals and the confusion of suchsignals. In some measurements, the chest acceleration signal as pickedup by the accelerometer 10 in FIGS. 1-3 had a rather slow varying, butvery strong (20-50 my peak-to-peak) motion component. Riding on top ofthis motion signal, was a higher frequency, but weaker (5-10 mypeak-to-peak) heart sound signal. Significant variability betweensubjects was observed in the frequency content of both the motion andthe heart sounds. Also, the two signals—motion and heart sounds—are notentirely frequency separable. Thus, simple digital band pass filteringdoes not consistently work to separate them. Physical motion impulsesfrom the feet couple very differently and in a non-stationary andnon-correlated manner to sensors placed at different parts of the bodyand also to orthogonal axes of the same sensor. Accordingly, even usingmultiple sensors to cancel out an artifact is complicated or unreliable.

Some of the embodiments remarkably introduce a Data Acquisition/SignalProcessing unit 20 with a special smoothing filter 130 in FIGS. 3-4 thattracks slow varying body motion signal wander or variation and thenremoves the wander from the sensor-based signal to give a clean(motion-removed) biomedical signal of interest on a Display unit 30. Thesmoothing filter 130 involves a polynomial filter or comparablyeffective smoothing filter used directly or in a composite signalprocessing path. Some embodiments use a subtraction step 140 as in FIG.4 to remove non-stationary motion artifacts reliably and robustly.Removing such artifacts makes the system more fully immune to sensorplacement and contact variations on the chest that might arise whenusing sensor 10. This provides a simple, yet effective way to reduce theimpact of motion artifacts and allow the reliable detection of primaryheart sounds and subsequent derivation of heart rate even when a personis walking while being monitored. In this way, motion signal removal orseparation, and heart-sound signal detection and heart-rate detectionare facilitated. No secondary reference or noise source is needed, thusreducing complexity of system design. Embodiments of structure andmethod thus extract primary heart sound signals from chest-worn sensor(e.g., accelerometer) data in the presence of motion artifacts.

Results from six subjects showed a primary heart signal detection rateof 99.36% with a false positive rate of 1.3% as described elsewhereherein (TABLE 2). Such type of embodiment appears to outperform noiseremoval techniques such as wavelet de-noising and adaptive filtering.(In certain motion conditions, or in combination, alternative approacheslike Wavelet Decomposition, Adaptive Filtering, Blind Source Separationmay in some embodiments also be used instead of, separately from,parallel to, or in combination with, the polynomial filtering.)

Advantages include: 1) uses as few as a single sensor or signal capturecomponent, 2) eliminates use of a secondary reference sensor, 3) allowsunobtrusive and non-invasive monitoring of vital biomedical signals inambulatory settings for continuous monitoring applications, 4) separatesheart signals independent of non-stationary bodily motion wander.

For biomedical instrumentation and signal processing for heart soundsspecifically, problematic motion artifacts are thus removed frombiomedical signals—such as from chest accelerometer signals and/or fromelectrocardiogram (ECG) signals—for use in ambulatory health monitoringsettings. The embodiments can also be extended by use of a spectrumanalyzer (Fourier analysis) to extract frequency separable components ofinterest too.

Ambulatory monitoring of cardiac activity can find widespreadapplications in home health monitoring of patients with a history ofcardiovascular conditions, monitoring older adults, ICU and hospitalmonitoring, monitoring vital signs in mobile ambulances, at accident andtrauma sites and can be used for fitness monitoring at exercise centersand elsewhere.

In some structure and process embodiments for removal of motion-relatedartifacts from biomedical signals, beneficial monitoring is providedfor, e.g., either or both of two independent signalsources—accelerometer 10 and ECG of FIG. 2. Chest acceleration signalsare collected in an ambulatory (walking) setting from real humansubjects using a chest-worn accelerometer 10—providing primary heartsounds signified as S1 and S2. Heart sound S1 includes audible soundsconcurrent with tricuspid and mitral valve activity and shows on aseismocardiogram as a pulse bundle. Heart sound S2 is mainly associatedwith pulmonary valve and aortic valve activity. Structures and processesof the embodiments thus remove motion artifacts and facilitate the useof a single, miniature, chest-worn MEMS accelerometer to pick up heartactivity and heart rate—derived from heart sounds—from ambulatorysubjects as shown in FIGS. 1 and 2. In FIG. 2, electrocardiogram signalsare independently collected from the human subjects walking briskly orrunning on a treadmill—providing signal components such as QRS from theECG.

Some background heart anatomy terms are as follows. De-oxygenated bloodenters right atrium of heart via inferior vena cava and superior venacava from systemic veins. The right ventricle of heart receivesde-oxygenated blood from right atrium and pumps it via the pulmonaryartery to the lungs where carbon dioxide is released and oxygen isreceived into the blood. The blood moves from the lungs via thepulmonary vein to the left atrium of the heart. Valves open and close atthe entry to, between, and exit from, the atria and ventricles. The leftatrium passes oxygenated blood to the left ventricle, which pumps theoxygenated blood out the large artery called the aorta. The aortaconnects by systemic arteries to cerebral, coronary, renal, visceral(splanchnic), and skin vasculatures and to vasculature of skeletalmuscles. The names of the valves are: tricuspid valve—right atrium toright ventricle; pulmonary valve—right ventricle to pulmonary artery;mitral valve—left atrium to left ventricle; and aortic valve—leftventricle to aorta.

The primary heart sound components, S1 and S2, are composite signalsgenerated by valve closures. S1 is caused by the closure of the mitraland tricuspid values of the heart, and S2 is caused by the closing ofthe aortic and pulmonary valves. An analog electrical heart monitoringsignal is captured by two or more ECG electrodes, and the signal is avarying voltage representing electrical activity of the heart, i.e., thesignal generated in a person's body to cause the heart to contract orrelax. The ECG signal has three main components, a P-wave, a QRS complexmade up of a Q-wave, an R-wave, and an S-wave, and a T-wave. The pulsesinclude a small positive P pulse, a larger negative-going QRSdepolarization pulse near in time to the S1 heart sound, and a largepositive-going T pulse near in time to the S2 heart sound. The P-waverepresents the depolarization (electrical activation) of the atria ofthe heart. The QRS complex represents the ventricular activity of theheart. The T-wave represents the re-polarization of the ventricles.

Process and structure embodiments can also be extended to otherbiomedical signals corrupted by motion wander—e.g., ECGelectrocardiogram, PPG—photoplethysmogram (signal from a PulseOximeter), EEG—electroencephalogram, EMG—electromyogram, ICG—ImpedanceCardiogram signals—or almost any other signal that might be affected bya separable wander. Thus, motion-related artifacts are removed from suchother biomedical signals in products that can be produced by amanufacturer in volume.

Remarkably, with some of the embodiments of structure and process,polynomial smoothing and differentiating functions and operations areperformed. A secondary reference sensor or signal source is unnecessary.Gross motion is tracked and canceled out from the primaryaccelerometer-based signal. A polynomial smoothing filter 130 (forexample, a Savitzky-Golay filter) is electronically instantiated hereinand digitally smoothes a given accelerometer-based data signal stream byapproximating it within a specified data window by a polynomial of aspecified order that best matches the data in the window in aleast-squares sense. Here, the electronic smoothing filter 130 fits theslower variations in body-motion-induced components of the biomedicalsensor-based signal and subtracts them as smoothed content from thebiomedical sensor-based signal to leave behind what is called a residuesignal. The residue signal provides a thus-extracted, faster-varyingsignal—primarily the heart sounds and other cardiac activity, as well assome residual or remaining noise.

Such polynomial filtering 130 preserves higher order moments aroundinflection points, or at extrema like peaks and troughs, that a digitalmoving average or low-pass filter does not. In other words, thepolynomial filtering better preserves features—like local maxima andminima—through a least-squares polynomial fit around each point. Also,unlike a moving average, in estimating the value of the fit at a certainpoint, it does not factor in the values on the polynomial fit around it,therefore not introducing a bias at such features while reducing thenoise.

In FIG. 1, a system embodiment has hardware that provides a measurementset-up and monitoring embodiment. A miniature (weight—0.08 gram,size—5×5×1.6 mm) triple axis, low-power, analog output MEMSaccelerometer (LIS3L02AL, STMicroelectronics, Geneva, Switzerland) istaped onto the chest (e.g., a few inches to the left of the sternumalong the third or fourth rib). (Taping the accelerometer sensor orusing a chest band presses the accelerometer sensor to or against a bareor shaved portion of the chest and efficiently couples chestacceleration to the sensor.) An acceleration signal corresponding to thecardiac activity is captured along the Z-axis—the dorso-ventraldirection orthogonal to the plane of the chest. The chest accelerationsignal is AC coupled with a 3 Hz cut-off and amplified with a gain of100 and low pass filtered—for anti-aliasing—through a three-stage,5-pole Sallen-and-Key Butterworth filter with a 1 kHz corner frequency.A commercial quad operational amplifier (op amp) package (LT1014CN,Linear Technology, Milpitas, Calif.) is used for the analog front-end.The accelerometer signal is then sampled at 10,000 Samples/sec using adata acquisition card (National Instruments, Austin, Tex.) and capturedand stored on a computer using MATLAB software (Version 2007b, TheMathworks, Natick, Mass.).

The AC coupling with approximately 3 Hz cutoff, which is a non-criticalrolloff frequency, is provided, for example, by a series couplingcapacitor C coupled to an input resistance established for theamplifier.

In FIG. 2, a reference ECG (lead II) is acquired simultaneously in athree electrode (single lead) electrocardiogram ECG amplifierconfiguration as a standard of reference in order to compare with theaccelerometer-derived cardiac signal for the evaluation of theperformances of the heart rate extraction from the accelerometer signal.

In FIGS. 3 and 4, for detection of primary heart sounds and cardiacactivity, the acceleration signal is digitally low pass filtered in astep 110 at 50 Hz—using a 3326 tap digital FIR filter with a steep 80 dBroll-off over 20 Hz—and decimated in a step 120 by a factor of 10.(Rolloff frequency less than 60 Hertz attenuates 60 cycle USA power lineinterference with biomedical signals of interest, and rolloff may bemade less than 50 Hertz for applicable countries using 50 Hertz. Whilethe rolloff frequency could be made higher, this FIR filter alsodesirably attenuates white noise above the frequency range of thesignals being monitored.) Also in a Phase 1, a high order Savitzky-Golaypolynomial smoothing filter 130, using 28th order and 401 point frame,is used to capture the relatively slow-varying motion wander and leaveout the more rapidly varying heart sound signal components. (Matlabsyntax for such filter is g=sgolayfilt(X,28,401) where g is the filteroutput and X is a latest input column vector of 401 sample values ofwindowed data.) In a Phase 2, the smoothing filter 130 output issubtracted in a step 140 from the decimated LPF output to obtain heartsounds S1 and S2. A folded correlation process in a step 160 thenenhances and strengthens the polynomial filtered S1/S2 peaks in themotion-removed acceleration signal. Such folded correlation process 160is described in further detail elsewhere herein and with background inU.S. patent application Publication “Heart Rate Detection In High NoiseConditions” 20100094150, dated Apr. 15, 2010 (TI-66732), which isincorporated herein by reference. Then the location of the peaks isthreshold-detected in a step 170 using an electronic amplitude-basedpeak picking process, and the selected peaks are counted in a step 180to calculate heart rate HR.

In FIG. 5, a chest-acceleration signal is derived from the accelerometersensor while a subject is walking around a room and low pass filtered at50 Hz (step 110) as shown in a first waveform. LPFing (low passfiltering) sub-50 Hz is used in some of the examples because most of thedesired signal power lies in that range and in general LPFing with somerolloff frequency below about one hundred Hertz in many of theembodiments avoids making the bandwidth so wide as to encompass andintegrate a substantial or undue amount of sensor noise (thermal, whitespectrum). In case LPF with a rolloff frequency above power-linefrequency is used, then some embodiments also include notch-filteringfor power-line frequency rejection. In FIG. 5, a second waveform is anelectronically-derived polynomial smoothing filter 130 outputcorresponding primarily to the body motion. A third waveformconcurrently shows the residue signal after subtraction 140 in FIG. 4and isolates the primary heart sounds. A simultaneous ECG timing signalis shown as a fourth concurrent waveform for reference.

In FIG. 6, the same embodiment monitors a chest-acceleration signal fromthe accelerometer Z-axis sensor while the subject walks on a treadmill.A brief rest recording is followed by motion. Compared to FIG. 5, theplot of FIG. 6 analogously shows an unfiltered (raw) acceleration signaland the residue from step 140 after the polynomial smoothing of step130. Note the magnified scale in some parts of FIG. 6, and that FIG. 6has a different scale than in FIG. 5.

FIGS. 7A-7C show signal plots for an ECG filtering embodiment 2. Theplots have different time scales and walking conditions. Raw ECG signalfrom the ECG electrodes in FIG. 2 and a concurrent filtered ECG signalwaveform, by applying steps 110-140 separately to the ECG signal, aredepicted for a subject walking on a treadmill.

In another embodiment, satisfactory S1-S2 heart signals were extractedfrom raw motion-affected accelerometer Z-axis data by LPF (low passfiltering) with corner at 100 Hz and then Savitzky-Golay filtering at20^(th) order, followed by subtraction of the S-G signal from the LPFsignal, and followed further by signal enhancement. It appears thatpolynomial filtering of motion-affected LPF accelerometer signals, usingpolynomial filtering on the order in a range of approximately 20^(th)order or higher order to at least over 30^(th) order, is satisfactoryfor obtaining heart signals as a residue by subtraction of thepolynomial filtering output from the LPF signals. Using polynomial fitsat such orders successfully captures both coarser and finer motioneffects. The smoothing filter in some embodiments can be lower order aswell, and may obtain good results even with a 1^(st)-order polynomial incase of some window sizes and applications. Also, lower order polynomialfiltering is contemplated and found useful as discussed laterhereinbelow. Using a number of points at least approximately half again(1.5 or more times) an order of the polynomial and even substantiallyhigher than that, in some of the embodiments, is believed to help toreduce noise.

In FIG. 8, a wireless embodiment has the accelerometer sensor 210 in achest-worn miniature unit including Bluetooth or pico-network wirelessor an RF transponder. The miniature unit 210 wirelessly communicateswith a Data acquisition/signal processing unit 215, 220 of FIGS. 8 and 4such as provided on a belt clip, in a cell phone or in a gatewayelsewhere in a residence (see FIG. 39). In FIG. 8, the signal processingunit 220 is coupled to a wireless modem 230 or transmitter (or to awireline modem) for transmission to a remote location such as a medicalclinic. The medical clinic has a receiver or transceiver such as in acell phone or wireless or wireline modem 240, and further has a datastorage and display unit 250. The medical clinic can interrogate theresidential data acquisition/signal processing unit 220 by transceiver240 via residence-based modem 230 and re-configure the residential unit215, 220 for various performances and for more or less information andmore or less frequent communications.

In FIGS. 3 and 8, any two or more, or all, of the described componentscan combined in a single digital system. The monitoring signal capturecomponent 210 is configured to capture a heart monitoring signal from aperson and provide it to an analog signal conditioning and samplingsection 215 (A-to-D) that feeds digital data to the signal processingcomponent 220. In some forms, the A-to-D happens physically within theaccelerometer chip and the signal flow remains electronically arrangedas shown. Either or both of components 210 and 215, 220 may provideamplification and noise reduction of the analog and/or digital signal inthe process. In various embodiments, the heart monitoring signal may beprovided to the processing component 220 in real-time, and/or may beprovided periodically as the signal is being captured, and/or may berecorded and provided to the processing component at a later time.

The digital heart monitoring signal may be provided to the dataacquisition 215 and signal processing unit 220 by wired or wirelessforms of communication, e.g., wired using a USB port, electrode wires,logic circuitry, etc. or wirelessly such as by a Bluetooth connection,Zigbee, or otherwise. In FIG. 8 a communications network for remotetransmission can be wide area network (WAN) such as the Internet, awireless network, a local area network (LAN), or a combination ofnetworks. Similarly, the processing component 220 may be connected to anoutput component 250 by any of the foregoing connections and networks.Any suitable display device and/or recording apparatus 250 is used suchas, for example, a computer monitor, a display of a handheld computingdevice, a display in a personal heart rate monitoring device, a displayin a piece of exercise equipment, etc. The system hardware of FIG. 39may be applied with one program at both the premises at which theaccelerometer is used and a replica of that system hardware applied withthat program and/or an additional program at the remote premises such asa medical clinic.

The system components including signal processing component 220 may alsobe implemented by or as part of any suitable digital system (e.g., ageneral purpose processor, a digital processor, a personal heart ratemonitoring system, a heart rate monitoring system in a piece of exerciseequipment, a personal computer, a laptop computer, a server, amainframe, a personal digital assistant, a television, a cellulartelephone, an iPod, an MP3 player, etc.) configured to receive thedigital heart monitoring signal from the monitoring signal capturecomponent 210. The processing component 220 is configured to process thedigital heart monitoring samples in the digital heart monitoring signalin accordance with embodiments of methods described herein. In one ormore embodiments of the invention, the processing component 220 includesfunctionality, e.g., a computer readable medium such as memory, a flashmemory, an optical storage device, a disk drive, flash drive, etc., tostore executable instructions implementing an embodiment of a method forprocessing heart monitoring samples as described herein and to executethose instructions.

Embodiments like those of FIGS. 8 and 4 and other Figures herein havepotentially wide-ranging applications from commercial products thatalready have in-built accelerometers (e.g., mobile phones, personalentertainment devices, content players, computer game controllers etc.)and those that do not (clothing, accessories etc.) to fitness products(heart straps, belts, wearable adhesive bandages or sensor tapes, clips,straps, bands or carriers for temporary affixing to one's chest, arm orelsewhere on the body, or implantable sensor devices) Embodiments aresuitably made as a part or whole of ambulatory monitoring products forambulances, at trauma sites (e.g., for accident or burn victims), forhome-monitoring of older adults and all populations to which theadvantages of the embodiments commend themselves.

The accelerometer 210 signals from all three axes are suitably alsoprocessed to electronically double-integrate the acceleration todetermine the location of the person wearing it. Since the person islikely to have been in bed overnight, the processing determines thelocation of the person during the day by double-integrating theacceleration starting from initial conditions of position initially atthe bed location, and zero initial vector velocity. This information canbe helpful as a cue to the person who is visually impaired, tocare-giver, and to a family member. The accelerometer processing canindicate that the person is in a given room of the residence, as anassist for one who is visually impaired, or can indicate that the personis leaving or has left the residence to inform a care-giver or familymember. In this way, the accelerometer and associated processor providenumerous services for all concerned, in various ways as taught herein.

For background on accelerometer calibration and double-integration seeU.S. patent application “Parameter Estimation for Accelerometers,Processes, Circuits, Devices and Systems” Ser. No. 12/398,775 (TI-65353)filed Mar. 5, 2009, which is incorporated herein by reference in itsentirety.

Due to its low-cost and ease of use, products using the embodiments havepotential for commercial success not only in urban and developed areasbut also widely in the developing world as well as in rural parts of thedeveloped world or in any place where low-cost, remote health monitoringfacilities may be rare, if available at all.

The smoothing filter 130 of FIG. 4 is configured based on a specifiedorder M and frame size (number of sample points N). For instance, aSavitzky-Golay polynomial smoothing filter is used in some embodimentsto best approximate the acceleration signal in the least-squares senseto capture the motion-dependent baseline-wander. In some embodiments,the smoothing filter is implemented in flash memory of a local processorof FIG. 39 such as a belt-worn unit or provided in a home networkgateway or clinic office network gateway, or cell phone or otherwise.

The matter of selecting and or finding feasible and optimum values fororder M and window length (N_(W) in points, t_(W) in time) for thepolynomial smoothing filtering is discussed next. In general for a fixedwindow length, N_(W), a higher order polynomial will fit the highfrequency components of the streaming data better. For a given order M,a shorter window of time will allow fitting the high frequency componentbetter.

In FIG. 4, the working hypothesis is that the accelerometer signal has alow-frequency (motion) component and a high frequency (heart signal)component. The polynomial filter is used to fit to the motion component.

A way to approach the optimization problem estimates the inherent orderof the low-frequency component and picks the smallest window thatsatisfies the condition that N_(W)>M+1 and N_(W) is odd (i.e.,N_(W)=2N+1). The smaller window size N_(W) is, the smaller is the numberof taps of the multiply-accumulate filter process implementing thesmoothing filtering. For an accelerometer signal in some applications,order M=1 and window size N_(W)=3 (sampling frequency is 1000 Hz). Insome examples herein, higher orders M and window widths N_(W) are shown.

In FIG. 4, motion, heart sounds and heart rate are electronicallyseparated and ascertained from accelerometer 210 data using thefollowing steps:

a) Low-pass filtering 110 and decimating 120 the accelerometer datab) Savitzky-Golay filtering 130 to fit the relatively lower frequencymotion datac) Subtracting 140 the output of the Savitzky-Golay filter from thelow-pass filtered accelerometer data (from step a) to obtain the heartsoundsd) Performing 160 folded correlation to enhance the primary heart sounds(S1 and S2) peak locationse) Peak picking 170 to count the number of S1 peaks in a predeterminedor configured segment (time interval) and counting 180 the heart rate HRin beats per minute BPM.

Note that the term ‘decimation’ refers to any process of regularlyremoving samples from a sample stream, or passing one sample in everyn_(D) samples as decimation parameter, and can but does not necessarilyrefer to removing all but 1 sample in ten. Thus, if a sample/ADCdelivers f_(S) samples per second, then a decimation process delivers adecimation frequency substantially f_(S)/n_(D) samples per second. If awindow period is t_(W) seconds, then the number of points N_(W)=2N+1 inthe window is N_(W)=1+f_(S) t_(W)/n_(D). The window period t_(W) may beselected by considering the time period over which the particularfeatures and behavior of interest are to be obtained by the filteringfrom the signal. The sampling frequency f_(S) may be selected with cost,physical size and complexity of anti-aliasing in mind (low pass filterAAF at 0.5f_(S) or less situated ahead of sampling f_(S)). The samplingfrequency f_(S) may be set substantially greater than the Nyquistfrequency for sampling the AAF output. The decimation parameter n_(D) isthen selected, firstly, to yield a decimation frequency f_(S)/n_(D) thatis sufficiently high relative to the e.g., 50 Hz low pass filter LPFfollowing the sampling/ADC circuit to provide effective operation ofthat LPF. Secondly, the decimation parameter n_(D) is also selected toyield a number N_(W) of window points that is sufficiently high relativeto the selected order M of the filter to keep filter noise low whilehaving the N_(W) window points being sufficiently low in number as tointroduce only so many filter computations as needed to achievesatisfactory filtering of the signal stream in the window. The filtercomputations are related to the product of the number N_(W) of pointsper window multiplied by a rate number r_(W) of windows processed persecond. If r_(W)=N_(W)/t_(W), the computations are proportional to N_(W)²/t_(W), which may motivate fewer window points and longer window timesin some energy-saving and lower cost processor applications. Remarkably,the examples herein satisfy these considerations for some applicationsand other examples may readily be devised for other particularapplications as well.

Mathematically expressed processes are described in further detail belowfor preparing various electronic embodiments with smoothing filters forvarious ways of motion extraction in step 130 and any other purpose towhich their advantages commend their use. They are appropriatelypartitioned into offline and real-time online electronic processes insuch embodiments.

The notation ∥(x−Ab)∥ in Equation (1) signifies the sum of squareddifferences between the [(2N+1)×1] respective data stream vector samplepoints or stream components and the (2N+1) respective estimates of thosestream components provided by multiplying a [(2N+1)×M] transform matrixA times a [M×1] vector of transform coefficients b_(j). The number oftransform coefficients b_(j) is M, and they form a [M×1] vector b. Agradient V is the [M×1] vector of first partial derivatives with respectto the transform coefficients b_(j). The number M of coefficients b_(j)is called the order, and if the number of transform coefficients b_(j)is M, then the order of the process is M. The [M×M] matrix of secondpartial derivatives with respect to the transform coefficients b_(j) issignified by ∇∇. The filter procedure involves, and in effect forms, acoefficients change coefficient vector Δb for updating an initialtransform coefficient estimate b=0 (i.e., all coefficients initializedto zero). This procedure pre-multiplies the matrix of second partialderivatives times the negative of the gradient to obtain that transformcoefficients change vector Δb.

Δb=−(∇∇∥(X−Ab)∥)⁻¹∇∥(X−Ab)∥  (1)

Since the Equation (1) involves a quadratic expression and starts fromb=0, the process directly finds the values of the transform coefficientsb=Δb in one pass without iterating additionally. Equation (2) representsthe result of performing the calculus operations represented by Equation(1). (Some embodiments transmit the coefficients b from Equation (2) toa remote site for record storage and further analysis, since theyeffectively compress much of the information in the data window. Ifcoefficients are to be transmitted, the [M×(2N+1)] matrix (A^(T)A)⁻¹A^(T) is pre-computed and then multiplied by each data window locally onthe fly. Other embodiments omit such compression and/or transmission, oronly do it locally on remote command, and thereby save some power andprocessing complexity.)

b=(A ^(T) A)⁻¹ A ^(T) X  (2)

This process generally finds transform coefficients b_(j) provided theinverse (A^(T)A)⁻¹ exists. That inverse exists when the rows of thematrix A are linearly independent (full rank) and enough data pointsN_(W)=(2N+1) are provided so that the corresponding number of columns ofthe matrix is sufficient for an inverse to be delivered.

In the special case of a polynomial transform process, a matrix ofindices is raised to powers, wherein the j^(th) column element A_(nj) inthe nth row of transform matrix A is raised to a power: n^(j). In otherwords, for the 2N+1 different values of n from −N to +N in the window ofa data stream X(i+n), the transform finds a set of coefficients b_(j)for a well-fitting power series to approximate all the values. Such apower series in general is represented by Equation (3):

X′(i+n)=b _(o) +b ₁ n+b ₂ n ² +b ₃ n ³ +b ₄ n ⁴ + . . . b _(M) n^(M)  (3)

Savitzky-Golay filtering outputs as the filter output g(i) for thewindow indexed i the value of b₀ estimated by Equation (2) for each datawindow, and successively window-by-window for successive indices g(i).

Rows of matrix A are orthogonal when the inner product is zero for anypair of different ones of them. These rows are illustrated in TABLE 1.The rows of values A_(nj) in matrix row n are non-orthogonal for theexample of a polynomial transform. (“̂A” signifies raising to a power.)

TABLE 1 ARRANGEMENT OF MATRIX A^(T) Power m (Order M) Points 0 1 2 3 . .. M n = −N: [1 (−N) (−N){circumflex over ( )}2 (−N){circumflex over( )}3 . . . (−N){circumflex over ( )}M]. . . . n = −4: [1 −4(−4){circumflex over ( )}2 (−4){circumflex over ( )}3 . . .(−4){circumflex over ( )}M] n = −3: [1 −3 (−3){circumflex over ( )}2(−3){circumflex over ( )}3 . . . (−3){circumflex over ( )}M] n = −2: [1−2 (−2){circumflex over ( )}2 (−2){circumflex over ( )}3 . . .(−2){circumflex over ( )}M] n = −1: [1 −1 1 −1   . . . (−1){circumflexover ( )}M] n = 0: [1 0 0 0 . . . 0{circumflex over ( )}M] n = 1: [1 1 11 . . . 1{circumflex over ( )}M] n = 2: [1 2 2{circumflex over ( )}22{circumflex over ( )}3 . . . 2{circumflex over ( )}M] n = 3: [1 33{circumflex over ( )}2 3{circumflex over ( )}3 . . . 3{circumflex over( )}M] n = 4: [1 4 4{circumflex over ( )}2 4{circumflex over ( )}3 . . .4{circumflex over ( )}M] . . . n = +N: [1 N N{circumflex over ( )}2N{circumflex over ( )}3 . . . N{circumflex over ( )}M].

Next, the process finds an estimated data stream X′=Ab.

X′=A(A ^(T) A)⁻¹ A ^(T) X  (4)

An electronic process is set up in a processing circuit as representedby Equation (2) and electronically executed by the processing circuit.For Savitzky-Golay filtering, the process is optimized to only find g(i)as the estimated value of b₀ and also to perform as much off-linepre-computation as possible. Accordingly, Equation (4) is revised as inEquation (5) to use only the n=0 row [1×M] of the first pre-multiplyingmatrix A instead of the whole matrix A in Equation (4),

g(i)=[1 0 0 . . . 0](A ^(T) A)⁻¹ A ^(T) X(i)  (5)

Sometimes a mathematical presentation of Savitzky-Golay filteringregards the window as multiply-added by a set of (2N+1) filtercoefficients c(n). Here, a [1×(2N+1)] filter coefficient vector C isintroduced so that

$\begin{matrix}{{g(i)} = {{C\; {X(i)}} = {\sum\limits_{n = {- N}}^{+ N}{{c(n)}{x\left( {i + n} \right)}}}}} & (6)\end{matrix}$

where

C=[1 0 0 . . . 0](A ^(T) A)⁻¹ A ^(T)  (7)

In Equation (8), an alternative notation CI equivalent to Equation (6)post-multiplies Equation (7) by a [(2N+1)×(2N+1)] identity matrix I anddesignates each of the columns of that identity matrix I as [(2N+1)×1]unit vectors E_(n). The phrase ‘unit vector’ for ε_(n) means a[(2N+1)×1] vector of all zeroes except for a one (1) at the nth rowposition. Furthermore, only the matrix inversion computations to formthe first row of inverse matrix (A^(T)A)⁻¹ are relevant and areperformed, considering the pre-multiplication by [1×M] row n=0 vector [10 0 . . . 0]. Thus, the filter coefficients are also equivalentlyexpressed in the notation of Equation (8), which is equivalent toEquation (7).

c(n)={(A ^(T) A)⁻¹(A ^(T)ε_(n))}₀  (8)

The Savitzky-Golay filter does a local polynomial fit in a least squaresense. For a given input variable data window x(i+n) and window oflength 2N+1 and chosen polynomial degree M, the filter output is givenby g(i). Filter coefficients c(n)—2N+1 of them—are computed, e.g.off-line, by electronic operations represented by Equation (7) or (8)and loaded into flash memory of a small signal processing unit eitherworn on the person or provided nearby and coupled wirelessly to theaccelerometer sensor 210 according to the blocks shown in the FIGS. 1-5.The signal processing unit suitably has a digital signal processorcircuit such as processor 220 that electronically performsmultiply-accumulates (MACs) represented by Equation (6) according to astored program accessing the filter coefficients c(n).

Some other embodiments use windows that are not centered around thevalue at index n used as the output (e.g., n=0). It should also beapparent from the above process description that a variety of choices ofmatrices A are possible and may be used instead of the particularpolynomial transform matrix shown in TABLE 1. The skilled worker choosesthe desired transform, the window (frame) size (e.g., 2N+1) and theorder M. Also, note that g(i) output of a first filter procedureproduces a data stream that itself can be windowed as represented bycolumn vector g₁(i₂) in Equation (9B). Accordingly, some embodimentsrepresented by Equations (9A), (9B) cascade two lower order filters ofEquation (4) and use straightforward technique to minimize theelectronic processing complexity of the computations in implementation.The transform matrices A1 and A2 can be the same or different, thewindow sizes (2N₁+1) and (2N₂+1) can be the same or different, and theorders M₁ and M₂ can be the same or different, all these choices beingindependent of each other.

g ₁(i ₁)=[1 0 0 . . . 0](A ₁ ^(T) A ₁)⁻¹ A ₁ ^(T) X(i ₁)  (9A)

g ₂(i ₂)=[1 0 0 . . . 0](A ₂ ^(T) A ₂)⁻¹ A ₂ ^(T) g ₁(i ₂)  (9B)

Some embodiments may also apply to the SG process a diagonal weightingmatrix W which is all zeroes in a [(2N+1)×(2N+1)] matrix except forweights down the main diagonal. The weights can, for instance, be one atthe middle of the diagonal and diminish symmetrically in value fartherfrom the middle of the diagonal. The motivation is that it may not beimportant for all points in the window to be well-approximated accordingto unweighted least squares, especially in a filter that is providing adetermination of one coefficient as output g(i). In that case, Equation(1) is replaced by Equation (10), which represents that the squares areeach weighted in the sum of squares ∥(X−Ab)∥:

Δb=−(∇∇∥W(X−Ab)∥)⁻¹ ∇∥W(X−Ab)∥  (10)

Then the electronic process represented by Equation (5) for the outputinstead is:

g(i)=[1 0 0 . . . 0](A ^(T) WA)⁻¹ A ^(T) WX(i)  (11)

The selection of transform type and matrix A is fixed/predefined byconfiguration or determined semi-static manner in some embodiments.Dynamic configuration or selection of the matrix A or transform type orparameters of a given transform is contemplated in some otherembodiments herein that determine which is the best transform type,order, window size, amount of cascading, etc. to use and thendynamically performs processing and remote communication.

Some other embodiments store and average a set of values from thetransform output of Equation (4) from different windowed segments of thedata stream X. This approach, roughly speaking, performs several filtersin parallel and averages them in an offset manner. All the valuesrepresent a reconstructed value corresponding to a same instant of time(i+n)=t, and note that this approach not only uses the n=0 row toapproximate b₀ but also uses the transform approximations to the othercoefficients that are available from Equation (3). In other words, theresults of approximating the data stream using 2N₁+1 successive windowsare used by selecting only the particular points that represent a giveninstant of time t. The number of points 2N₁+1 averaged (say, some numberin a range 3 to 11 points) is enough to average out some noise withoutmuch extra computer burden N₁<=N. Those points are the successive windowdata X at indices n=t−i such that for succeeding windows i, theapproximate data values X′ generated by the power series start at highindex n=+N₁ and proceed down to n=−N1. The electronic processor 220(and/or 240) executes instructions or otherwise performs the electronicprocess as represented by Equation (12), where X′(i+n) is from Equation(4). Equation (12) reduces to Equation (6) when N₁=0 (i.e., 2N₁+1=1).

$\begin{matrix}{{g(t)} = {{\left\lbrack {1/\left( {{2N_{1}} + 1} \right)} \right\rbrack {\sum\limits_{n = {{- N}\; 1}}^{{+ N}\; 1}{X^{\prime}\left( {i + n} \right)}}}_{({{i + n} = t})}}} & (12)\end{matrix}$

In view of the analysis herein, it is emphasized that other types ofprocesses can be alternatively selected according to the teachingsherein, whether they are called Savitzky-Golay or not. The skilledworker sets up a test bench with library accelerometer-based waveformsand then makes the transform matrix choices, choice of number of points(2N+1), and choice of order-value M, either manually or by an automatedprocess. The filtering choices are tested either by visual inspection ofa display of output from FIG. 1 process or by automated processaccording to metrics of false negatives and false positives, etc. asdescribed herein. Transform matrix A values, and values of (2N+1) and M,2N₁+1, etc. for one or more such filter processes having favorablemetrics are then loaded into the monitoring device flash memory or harddrive and executed in real time on processor 220 (or 240).

A transform for an embodiment approximates an actual data stream vectorx(i+n) and produces an output signal stream that follows the heart soundpeaks well over time in response to a data stream X herein derived froma body-worn accelerometer. Some embodiments have reduced processingcomplexity by using low enough frame size (2N+1), order M and/or usingan efficient transform matrix A to achieve desired performance for thepurposes for which the monitoring is intended. The same transform isdesirably low-complexity and well-performing over numerous patients,accelerometers and their positioning on the body, and in differentenvironments of use, such as clinic, hospital, home, exercise venue,etc.

In FIG. 4 and turning to a succeeding electronic process portion 150 forenvelope-based noise rejection, an amplitude envelope is generated, asshown in FIG. 9. In FIG. 9, the residue signal stream r(i)=x(i)−g(i) ofFIG. 4 has some remaining noise from subtracting step 140 that subtractsthe smoothing filter output g(i) from the LPF-supplied input x(i). Anenvelope is fitted to the residue as indicated by the dottedenvelope-line in FIG. 9. The amplitude-based processed output R(i) isshown in FIG. 9, as derived from the envelope-fitted residue r(i). Thenoise n(i) near the horizontal axis is substantially reduced. Operationsfor this process suitably use an envelope-related variable gainfunction. Alternatively, the circuit and/or process is arranged togenerate zero signal output when the envelope is below a low thresholdthat still passes the peaks.

Description next turns to the FIG. 4 electronic process portion 160called folded correlation. For background, see the incorporated patentapplication publication TI-66732. An input data stream of residuer(i)=x(i)−g(i), or envelope-processed residue R(i) comes to the foldedcorrelation process. Recall that g(i) is the output of the smoothingfilter such as represented by Equation (6) or from a buffer memory forit. Processed filtered residue R(i) is windowed by a further data window(also called a frame) of length 2N₂+1 (with points accessed by an indexn=0, 1, 2 . . . N₂).

The output f_(c)(i) of the folded correlation is given by Equation (13):

$\begin{matrix}{{f_{c}(i)} = {\sum\limits_{n = 0}^{N\; 2}{{R\left( {i - N_{2} + n} \right)}{R\left( {i + N_{2} - n} \right)}}}} & (13)\end{matrix}$

The digital data stream for heart monitoring residue signal samples R(i)from the smoothing filter subtraction is successively processed inoverlapping frames indexed i. In general, the value of 2N₂+1 is selectedto be approximately the width t_(W) of a desired signal event (e.g., anS1, S2, or R-wave). For example, S1 is typically about 100-150milliseconds long. If the decimated sampling frequency f_(S) is 1000Samples/sec, the value of 2N₂+1 is established, e.g., as an odd numberbetween 101 and 151, and N2 is some number between 50 and 75 inclusive.Thus,

N2=RND(t _(W) f _(S)/2)  (14)

In some embodiments, the value of N2 is configured in flash memory, andcan be selected or altered by a local or remote operator of a FIG. 8remote digital system portion 240, 250 using the embodiment of structureand process.

In the electronic folded correlation process 160 represented by Equation(13), the heart monitoring residue samples R(i+N2−n) from the later halfof each frame are folded around the center heart monitoring sample R(i)in the frame and multiplied by dot product (sum of products in Eq. (13))with heart monitoring residue samples R(i−N2+n) in the earlier half ofeach frame. The result of the dot product is a folded correlation outputsignal stream f_(c)(i) corresponding to instant i of the input residuesignal stream R(i) in the center of the frame.

In FIG. 10, the residue signal input R(i) (subtraction of SG fit fromLPF input) is shown in a lower waveform and an output f_(c)(i) of FoldedCorrelation 160 is shown in an upper waveform. Sharp, distinct, positivepeaks in output signal f_(c)(i) are output by Folded Correlation 160.This is because not only positive peaks but also negative peaksfolded-correlate positively with themselves due to multiplication(++=+,−−=−). Between the peaks, the noise folded-correlates with itselfnegligibly. The resulting output f_(c)(i) as a whole recovers pulsesthat follow S1 and S2 heart sounds well.

Succeeding thresholding passes the S1 peaks and counts them. Robustdetection of primary heart sounds and heart rate from a chest-wornaccelerometer is thus achieved in the presence of interfering motionartifacts. Such capability is directly relevant in applications thatinvolve ambulatory monitoring of cardiovascular and cardio-respiratoryhealth. Applications include: home health monitoring, fitnessapplications (exercise monitoring), hospital and ICU (intensive careunit) patient monitoring, and patient monitoring at accident sites, inambulances, gurneys or rolling patient transfer beds, in mobility aidslike scooters and wheelchairs, and other mobile and/or fixedenvironments in a setting that is related to a hospital, clinic, alliedmedical testing facility, residence, commercial establishment, airportor otherwise.

Heart signal components may have S1, S2, and heart murmur components.Some embodiments further process heart signal components by coupling thecircuitry and signals described herein to processing according to theteachings of U.S. Patent Application Publication 20090192401 “Method andSystem for Heart Sound Identification” dated Jul. 30, 2009 (TI-65798),which is incorporated herein by reference.

Data used for the evaluation of the methods is collected from sixhealthy young volunteers. Ambulatory conditions were simulated by thesubject walking 2-3 minutes at normal speed.

Table 2 shows the accuracy, number of false positives and number offalse negatives for the subjects collectively. Most of the falsenegatives were due to S2 misses.

TABLE 2 ERROR ANALYSIS OF RESULTS Accuracy 99.36% False Positives  1.3%No. of S1 misses   2 (0.085%) No. of S2 misses 13 (0.55%)

FIG. 11 shows a Bland-Altman plot for heart rates (calculated over every5 second time segment) for all of the subjects from ECG data and heartrate estimates obtained from the processed accelerometer data. Thedifference of the two heart-rate measures (accelerometer and ECG) isplotted versus their average. A few outliers are caused by falsepositives, but overall, most of the data is within a 95% confidenceinterval. For heart rate calculation both S1 and S2 locations were used.The results are robust over both lower and higher heart rates.

In some embodiments, the odd peaks from the output of a simple form ofthe peak detector are picked as S1 and the even peaks as S2. This typeof selection may lead to some errors since a single false peak can causethe error to ripple along. The effect of this is mitigated to someextent by a choice of performance measures that look at relative timedisplacement or distance in time as opposed to absolute location intime. Nonetheless, the simple form of the peak detector and processlocates most of the S1 and S2 events with very few false positives. Insome other embodiments, to reduce S2 false negatives and reduce falsepositive rate even further, the peak detector is augmented with acircuit or process that incorporates amplitude and S1-S2 intervalinformation to select the S1 and S2 peaks from the output of the peakdetector.

In FIG. 12, to further illustrate the ability of a process embodiment tolocate S1 events robustly, the peak detection is made to pick only S1events. FIG. 12 shows a plot of Cardiac Interval from S1 locations fromaccelerometer data (S1-S1) on the graph vertical axis, versus ECG R-Rinterval on the graph horizontal axis. As seen from FIG. 12, a highcorrelation (correlation coefficient of 0.98) between the cardiacperiods was obtained from the different measures. The slope of the leastsquares fit is 0.99.

Benefits are obtained by themselves and with other benefits bystructures and processes described elsewhere herein and in thesimultaneously-filed TI-68552 and TI-68553 patent applications, whichare incorporated herein by reference.

Description turns next to a set of embodiments that separate and derivemotion/activity, heart-rate and respiration from a single signal from asingle chest-worn sensor such as a miniature Z-axis accelerometersensor. Ambulatory measurement of respiration and cardiac activity canfind wide application in home health monitoring of older adults and ofpatients with a history of cardiovascular, respiratory, and otherconditions for which respiratory and/or cardiac monitoring are desired.Evaluating cardiovascular performance of patients in ICU and hospitalsettings, in mobile ambulances, and at accident and trauma sites alsocalls for ambulatory cardiac and respiratory measurement and monitoring.Conventional solutions for heart-rate and respiration monitoring arebelieved to be expensive, invasive or obtrusive and too cumbersome forambulatory and continuous monitoring applications.

Remarkably, various embodiments with a single, miniature, chest-wornMEMS accelerometer and associated monitoring circuitry measure andmonitor respiration, motion and heart activity—reflected by heartsounds—as shown in FIG. 13.

In FIG. 13, embodiments are provided for ambulatory monitoring ofheart-rate and heart sounds, activity, body motion and respiration in anon-invasive and minimally obtrusive way. Here, a single sensor, such aswith a MEMS accelerometer, extracts not only heart-rate/heart sounds butalso respiration in an ambulatory setting. Any signal is used thatincludes detectable heart sound signals from at least one sensor axis ofthe accelerometer sensor, or from two or more sensor axes. When a bodymotion signal component is included in the sensor signal, that bodymotion signal component is in some embodiments separated out or isolatedand delivered as a useful output representing activity or motion aswell.

Some advantages of various embodiments are extraction of three vitals(respiration, activity, heart sounds/heart-rate) from a single sensorand a single signal. A miniature sensor embodiment taped on the chestprovides a non-invasive and minimally obtrusive way to sense and monitorvital physiological parameters in the presence of motion. Embodimentscan be used with minimal inconvenience in ambulatory and continuousmonitoring applications, and are very inexpensive and can be made intodisposable patches and tapes, for instance.

In FIG. 14, a respiration waveform is obtained, for example, by amonitoring device embodiment of FIG. 13 and its process embodiment. Theprocess in FIG. 14 receives the raw signal stream from the FIG. 13 ADC(analog to digital converter) and then first separates the heart soundsfrom the composite signal from the sensor using Savitzky-Golay (S-G)polynomial fitting followed by Folded Correlation and Peak Detection todeliver the S1 heart signal peaks. The heart rate is counted in responseto the peak detection to provide a Heart-Rate signal output.Concurrently, the respiration is monitored by then measuring thesuccessive times, called the inter-beat intervals or S1-S1 intervals,between heart beats—beat-by-beat. The variation in the measuredinter-beat interval over time thus represents respiration because it isrespiration-dependent and substantially independent of non-respiratorygross body motion. The monitoring device thus delivers as a respirationwaveform that substantially represents the inter-beat interval varyingover time. Further respiration processing counts the breathing rate anddelivers a resulting breathing rate output, and derives and outputs anyother useful information. In the meantime, the motion signal isextracted either from the S-G polynomial smoothing filter 130 as in FIG.4 or by a low pass filter LPF with corner frequency at 2 Hz as shown inFIG. 14.

In FIG. 14, post-processing of the motion signals is applied to monitorand deliver waveforms representing average activity level over time,monitor walking gait and other motions, and detect a fall if one were tooccur. Consequently, deriving motion or activity from an accelerometeris important, such as by the present embodiments. For instance, averageactivity level can be generated as the root-mean-square (RMS) of themotion waveform measured over an hour and output hour-by-hour. Walkinggait can be derived from the Z-axis alone or in combination with signalstreams from other accelerometer axes with respiration subtracted out. Afall is indicated such as by a peak detection of an unusuallyhigh-magnitude acceleration peak which stands out from any recent orsubsequent neighbor peaks in a predetermined window of time such as+/−15 seconds. Some embodiments thus deploy motion-based analysis as afall sensor as described herein or otherwise in any suitable mannerenabled by an embodiment.

Process embodiments as in FIGS. 3-4 separate motion signal componentsfrom an accelerometer sensor signal to cleanly and robustly extractheart sounds, which enables the use of the accelerometer 210 to monitorheart sounds in the presence of motion and activity. Moreover, processembodiments as in FIGS. 13-14 separate motion signal components from anaccelerometer sensor signal to cleanly and robustly extract heart soundsand use the accelerometer sensor to monitor not only heart sounds butalso respiration (derived from heart sounds) in the presence of motionand activity, and further to deliver a motion/activity signal as well.In some embodiments, motion-based gating is performed to reject signalframes in the event that the motion/activity level is unusually high anddoes not permit reliable detection of heart-rate or respiration or someother derived signal under a given detection process. Reliablecancelation of motion artifacts from accelerometer signals to extractheart sounds is described, among other things, hereinabove and in thesimultaneously-filed TI-68518 patent application, which is incorporatedherein by reference.

In FIGS. 13-14, one example of a monitoring system embodiment has thesame hardware and accelerometer sensor description as given inconnection with FIGS. 1-4. Respiration monitoring is added as in FIG.14. A reference ECG may be provided as discussed for FIG. 2.

In FIG. 15, detection of a body motion waveform is shown. The monitoringdevice measures gross body motion to sense and monitor activity and canprovide a useful index of a person's level of activity over a period oftime, and facilitate inferences about the person's lifestyle andmetabolic index—jointly with ECG-derived heart-rate. Gait recognition byaccelerometer-based motion monitor aids biometric assessment and canidentify signs of or precursors to a dangerous fall. Also, using theaccelerometer-based motion monitor as an indicator of sudden,high-magnitude acceleration can additionally identify signs of orprecursors to a dangerous fall, as well as a fall itself.

In FIGS. 15 and 14, the signal obtained during motion from thechest-worn accelerometer is digitally low-pass filtered at 2 Hz—using adigital FIR filter—to extract the slowly varying baseline wander due tomotion. For at least some cases of body motion as well as the body atrest, the respiration signal is well decoupled in frequency from thelow-frequency body motion signal and is thus digitally low-pass filteredto successfully extract the respiration signal. FIG. 15 shows the rawand filtered motion signal extracted from the accelerometer during restand brisk motion. Motion/Activity extraction from the accelerometer isconveniently achieved.

The implementer pays attention to physical attributes of the sensors inorder to reduce unnecessary or activity-irrelevant motion artifacts.Coupling noise is reduced through good sensor location and placement andsecure attachment of the sensor. Wire line noise or cable noise is keptlow or eliminated by intelligent selection and placement and secureelectrical and physical attachments. Wireless transmission isalternatively used to couple the hardware components of the monitoringsystem to reduce or eliminate body motion effects other than thosepicked up by the accelerometer itself and included in the accelerationsignal(s).

In FIG. 16, sensing, detecting, and monitoring heart rate, heart sounds,and cardiovascular and cardio-respiratory activity during motion andexercise importantly support continuous patient motoring and fitnessapplications, as well as at emergency and accident sites. Theaccelerometer-based extraction of primary heart sounds—S1 and S2produced by the heart valve pairs closing at the ends of the diastolicand systolic periods respectively of the cardiac cycle—is robust in theabsence or presence of motion, as shown in FIG. 16. The primary heartsounds are robustly detected through the S-G processing of the chestacceleration signal, not only during resting conditions, but also in thepresence of strongly interfering motion—like walking

In a respiration monitor example for FIG. 16, the S-G digital filtering130 and residue generation 140 by the processor to obtain the heartsounds are as described in connection with FIGS. 3 and 4. Timing- andamplitude-based thresholding 150 and Folded Correlation 160 are appliedas process steps of FIG. 4 as described earlier hereinabove. The FoldedCorrelation 160 in an example has a frame size of 7 at with the 1000samples/sec that resulted from decimation. The peaks in FIG. 10corresponding to S1 and S2 in the motion-removed acceleration signal arethus strengthened and are then peak-detected in step 170 for counting180.

FIG. 16 shows as a first waveform the raw acceleration signal low passfiltered at 50 Hz. A second waveform represents theelectronically-performed numerical polynomial fit g(i) correspondingprimarily to the motion. A third waveform is the residue r(i) aftersubtracting the second waveform from the first waveform. The thirdwaveform exposes or isolates the heart sounds, with motion-inducedamplitude variations of FIGS. 13 and 16 as amplitude modulation thereon.In FIG. 16, the timing locations of heart sound components S1 and S2show plainly and precisely, and they are largely independent of theamplitude modulation except for respiration-related variation ofinter-beat interval. A fourth waveform in FIG. 16 is the simultaneouslyacquired ECG signal that is used as a timing reference or referencestandard. Heart sound detection in the presence of motion is thusachieved.

Detection of respiration from the inter-beat interval has aphysiological basis. Respiration modulates the heart rate, andconsequently the inter-beat interval, by a phenomenon called respiratorysinus arrhythmia RSA, which is possibly responsive torespiration-related and other intrapleural or intra-thoracic pressurechanges. The respiration-dependent variation in inter-beat interval isconventionally obtained from the R-R interval in the ECG recording. R-Rinterval robustly tracks respiration even during motion and exercise.

FIGS. 17, 19, and 22 respectively show three processes or methods ofelectronically generating respiration signal outputs from a sensorinput. In various embodiments, these processes are used eitherindividually, or in pairs, or a combination of all.

Some embodiments as illustrated in FIGS. 13, 14, 16 and 17 make the ECGrecording optional, or obviate and eliminate the ECG recording, byderiving respiration from an accelerometer sensed-and-residue-detectedvariation in the S1-S1 interval during motion.

In FIG. 17, a process embodiment obtains the heart sound signal at astep 310 by removal of motion-dependent wander in Phase 2 of FIG. 4. Theheart sound peaks are reinforced through Folded Correlation 320, andpeak detection 330 is performed to detect the S1-S1 peaks. In a step340, the S1-to-S1 interval (or S2-to-S2 interval) is repeatedly computedbeat-by-beat to electronically obtain data values of successiveinter-beat intervals. These data points are interpolated in a step 350to yield a continuous respiration waveform of FIG. 18.

In FIG. 18, a first waveform shows the raw acceleration signal fed fromthe 50 Hz LPF to the smoothing filter in FIG. 4. A second waveform showsthe residue signal that delivers filtered heart sounds S1/S2. A thirdwaveform shows a varying inter-beat interval—the S1-S1Respiration-related waveform—derived according to the process of FIG. 17and superimposed on a concurrent ECG (R-R) derived respiration waveform.The fit is quite favorable, as disclosed by inspecting the two differentscales of illustration in FIG. 18.

In FIG. 19, another process embodiment is called the baseline wandermethod herein for respiration monitoring by a single accelerometersensor. This process in a step 380 operates on the raw accelerometer ADCsignal input from a person at rest by low-pass filtering it with afilter cutoff frequency at about 2 Hertz, or otherwise selected, e.g.,with LPF cutoff in a range of about one (1) Hertz to about three (3)Hertz. A waveform called Baseline Wander thus obtained as an electronicrespiration signal at a step 390, with example waveforms shown in FIGS.20 and 21 for comparison with respiration outputs from each of theprocesses of FIGS. 17, 19 and 22. In other words, breathing periods ofabout half a second or more are passed, so that not only restingbreathing periods of a breath every two or three seconds are detected,but also breathing periods under stress or after exercise down to abouta third or half a second are detected. Shorter period signal variationsin the accelerometer are suitably obtained from the S-G filter smoothingor by other types of filtering to represent body motions other thanrespiration.

In FIG. 20, a comparison of concurrent waveforms of respirationgenerated by different embodiments is shown along with a referencerespiration waveform obtained from a respiration belt or spirometer. Asecond waveform shows a baseline wander signal using the processembodiment of FIG. 19. A third waveform shows the varying inter-beatinterval method or process of FIG. 17 (RSA: S1-S1 interval), and afourth waveform shows output from a heart sound amplitude modulationprocess embodiment of FIG. 22.

In FIG. 21, another comparison of concurrent waveforms of respirationgenerated by different embodiments is shown. A first waveform shows thevariation of S1 amplitude output from a heart sound amplitude modulationprocess embodiment of FIG. 22. A second waveform shows a trace of S1power obtained by further processing the first waveform. A thirdwaveform shows the varying inter-beat interval method or process of FIG.17 (RSA: S1-S1 interval). A fourth waveform shows a baseline wandersignal using the process embodiment of FIG. 19. A fifth waveform shows areference respiration signal. A sixth waveform shows an ECG-derivedrespiration signal from inter-beat interval obtained from the R-Rinterval of the ECG. A seventh waveform shows a trace of filteredacceleration.

In FIG. 22, a further process embodiment, for respiration monitoring bya single accelerometer sensor 210, obtains the heart sound signal at astep 410 by removal of motion-dependent wander in Phase 2 of FIG. 4. Theheart sound peaks are reinforced through Folded Correlation 420, andpeak detection 430 is performed to detect the S1 peaks. These S1 peaksare amplitude modulated as shown by the second (middle) waveform of FIG.23. Instead of (or in some embodiments in addition to) inter-beatinterval measurement as in FIG. 17, the successive S1 peak amplitudes(or S2 peak amplitudes) in a step 440 are repeatedly measuredelectronically beat-by-beat in FIG. 22 to electronically obtain datavalues of successive heart sound peak amplitudes. These data values areinterpolated in a step 450, such as by linear or quadratic or otherinterpolation, to yield a continuous respiration waveform. Examples arecomparatively shown as the first and second waveforms of FIG. 21 foramplitude and power respectively.

FIG. 23 illustrates the close correspondence of respiration measurementsobtained in different ways. In a first waveform from ECG, amplitudemodulation rides on the R peaks. Amplitude modulation rides on the S1peaks in the second (middle) waveform from accelerometer based sensingaccording to the process embodiment of FIG. 22. The S1 amplitudemodulation correlates well, as seen by comparison with the R amplitudemodulation on the first waveform from ECG. A third waveform shows arespiration signal obtained from a respiration belt for reference.

FIG. 24 shows an implementation of a wired system embodiment 600 for arespiration and cardiac monitoring system. An accelerometer 510 isstrapped to the chest of the person being monitored. An axis sensorsignal is coupled to a data acquisition signal processing unit 520having a stream data interface and an associated data storage unit 530for the signal stream and for instructions and parameters. The signalprocessing unit 530 supplies process monitoring data to one or moredisplay units 550.i, each having a respective data storage unit 560.i. Afirst form of display 550.1 shows the heart sound signal and/or heartrate. A second form of display 550.2 shows the body motion signal. Athird form of display 550.3 shows the respiration signal and/orrespiration rate and/or or respiration depth (how deeply the person isbreathing) and/or other respiration parameters. Various parameters forrespiration are obtained from the respiration waveforms by findingvarious values on the waveforms and differences and trends therein. Forexample, respiration rate is measured as the number of cycles ofinhalation and exhalation in a given time window (e.g. one minute).Averaging and signal fusion methods/algorithms are also usable in someembodiments to derive more robust respiration rates from multipleparameters. For instance, how deeply the person is breathing isrepresented by an average of the difference between successive values ofinhalation peak and exhalation trough in a given time window (e.g. oneminute). Averages and trends in the inhalation peaks are readilycalculated and displayed. Averages and trends in the exhalation troughsare also readily calculated and displayed.

The system 500 of FIG. 24 is suitably arranged and physically protectedfor mobile and ambulatory monitoring environments. In other forms thesystem 500 is set up for use in a clinic by one or more cliniciansconcurrently.

FIG. 25 shows an implementation of a wireless system embodiment 600 fora respiration and cardiac monitoring system including various remarkabledevice or component embodiments. The description parallels that of FIG.24, except that the accelerometer sensor 610 and its electronic circuitalso have a Bluetooth or other short range wireless modem wirelesslycoupled to another short range wireless modem 622 that is coupled via astreaming data and control interface 624 to a data acquisition signalprocessing unit 620. Further, modems 640.2 and 670 for wirelesstransmission and reception remotely are provided at each of twolocations so that the data acquisition signal processing unit 620communicates via its modem 640.2 to the remote wireless transceiver unitor modem 670. The latter modem 670 is coupled to be one or more displayunits 650.i and their storage unit(s) 660.i. In this way, tele-medicineapplications are supported. The acquisition signal processing unit 620and its modem 640.2 are suitably provided in a residence or ambulance oron the person or in a wheelchair or gurney. The wireless transceiver 670and display unit(s) 650.i are suitably provided in a clinic, hospital,medical monitoring center or otherwise. Either or both ends of thewireless system may be mobile, such as one example of a modem 640.3 andalert/processor/display 680 when a professional in a vehicle is urgentlyneeded to review data coming in from a residence or another vehicle incase of emergency and to respond with instructions.

In FIG. 25, combinations with further processes, circuits and devicesfor automatic cautionary responses, warnings, and/or automated monitoredtherapeutic responses are contemplated. Upon occurrence of undueexcursions of one or more measured parameters or relationships amongparameters detected by signal processing unit 620, the remote processor670 alerts any one or more of medical professional, patient, caregiver,and/or family member via a modem 640.3 and alert/processor/display unit680 by sending a cellular telephone call and/or other voice call ormessage and/or written alert such as an automatic e-mail. The alertsystem suitably provides for acknowledgement by any of the recipients.Also, another modem unit 640.1 is suitably provided and coupled to atele-medicine therapeutic or assistive device 690 for assisting thepatient in some pharmacological, informational, or physically assistiveway by administering a medicine, or adjusting a dosage or otherwise. Incase of excursions that indicate an extreme medical emergency, the dataacquisition signal processing unit 620 may be permitted to locallycontrol the therapeutic or assistive device 690 temporarily and in amaximally-safe way until remote commands are received or firstresponders can arrive. Mere removal or inadvertent detachment of theaccelerometer 610 from the chest is distinguished by the electronicprocessing 620 from affirmatively detected excursions of measuredsignals and parameters. Regarding tele-care assistance, such assistanceis suitably rendered in some physical way in response to the real-timeaccelerometer sensor 620 data by activating motorized apparatuscomprehended by device 690 such as to adjust a motorized bed, or move ascooter into proximity for patient use, or servo-mechanically actuateand flush a toilet unit, or open a tub drain to empty a tub, or someother assistance.

In FIGS. 24 and 25, various parts of the systems 500 and 600 are eachvariously miniaturized and partitioned into various modules and providedwith various types of wireline interfaces or wireless modems fordifferent types of products. In this way, different system embodimentsare provided. One type of embodiment forms a complete medical clinicsystem. Another type of embodiment is a patient-worn medical-sensorand/or therapeutic device that is wired or has a wireless modem. Anothertype of embodiment is a patient-worn signal processing and modem moduleon a belt clip that connects or wirelessly couples to such a sensor andwirelessly couples to a premises gateway or directly transmits to aremote location. Another type of embodiment includes the sensor, signalprocessor, memory, and modem together in a micro-miniature device thatis taped to the chest and communicates to a router or gateway locally,and the latter communicates remotely. Another type of embodiment is thelocal router or gateway that includes signal processor, memory, andmultiple types of modem to communicate with the sensor and separatelycommunicate remotely, such as in a patient home-based system totelecommunicate to clinic or hospital. See FIG. 39 and FIGS. 40A/40B foran example of apparatus to support these various embodiments.

Respiration detection and monitoring for a person performing body motionor at rest are thus conveniently achieved along with cardiac monitoring.Local/remote assistance is suitably initiated responsively to suchdetection and monitoring. By contrast, conventional respirationmeasurement devices like respiration belts and spirometers are verysusceptible to motion-dependent artifacts and/or are very unwieldy forcontinuous ambulatory monitoring. Various embodiments can thereforesignificantly facilitate the measurement of respiration in the presenceof motion, or at rest. Benefits are obtained by themselves and withother benefits by structures and processes described elsewhere hereinand in the simultaneously-filed TI-68553 and TI-68518 patentapplications, which are incorporated herein by reference.

Description turns now to further embodiments for estimation of bloodflow and hemodynamic parameters from a single chest-worn sensor.Embodiments are provided for measurement of blood flow trends (strokevolume, cardiac output) and other hemodynamic parameters (contractility,pre-ejection period, iso-volumic contraction interval) in a non-invasiveand minimally obtrusive way. These measurements are believed to havebeen problematic, expensive, and inconvenient in the past. Conventionalhemodynamic monitoring e.g., some forms of Doppler echo or impedancecardiograms, and some blood pressure monitors, may be expensive,invasive or obtrusive and too cumbersome for ambulatory and continuousmonitoring applications. Here, by contrast, a single miniature sensorsuch as a MEMS accelerometer coupled with a data acquisition signalprocessing embodiment extracts hemodynamic parameters from the in-planevertical accelerometer axis (Y-axis). These benefits are obtained bythemselves and with the respiration detection and other features thatare described hereinabove and in the simultaneously-filed TI-68552 andTI-68518 patent applications, which are incorporated herein byreference.

Among the advantages of some of the present embodiments, are:

a. Uses a single sensor and a single signal to extract severalhemodynamic vitals such as any, some or all of changes in stroke volume,changes in cardiac output, heart-rate, isovolumic contraction interval,etc.b. Is minimally obtrusive (miniature sensor taped on the chest)c. Can be used with minimal inconvenience in continuous monitoringapplicationsd. Disposable patches/tapes carry the sensor and offer low cost andconvenience.

Embodiments of system, circuits and process enable the use a singlechest-worn miniature sensor (e.g., a dual-axis or triple-axisaccelerometer) for the extraction of a signal closely related to theflow of blood from the heart. This enables extraction and assessment ofother hemodynamic and cardiovascular parameters such as those discussedabove and in the next several paragraphs.

Isovolumic contraction interval IVCI is the duration of an event duringthe early systole when the heart ventricles contract without any changein volume. During the isovolumic contraction interval the myocardialmuscle fibers have begun to shorten but have not developed enoughpressure in the ventricles to overcome the aortic and pulmonaryend-diastolic pressures and thereby open the aortic and pulmonaryvalves. Such contraction interval occurs after the closure of the mitraland tricuspid valves and before the opening of the semi-lunar valves.Both pairs of heart valves are closed during this interval. IVCI can beestimated as the time difference between the peak of the S1 waveform(from the normal axis or Z-axis of the accelerometer) and beginning ofthe first peak of the vertical Y-axis accelerometer signal. Thisinterval ICVI is expected to correlate well with the time difference inFIG. 32 between the time P1 of occurrence of the S1 sound and the peaklocation in time F1 of the flow signal.

Stroke volume SV is the difference between the end diastolic volume andend systolic volume and is a measure of the blood pumped by the heartper cardiac cycle. A conventional pulse contour method calculates ablood flow variable (milliliters/sec) from the pressure signal andcomputes the stroke volume by integrating the blood flow signal over acardiac period. By embodiments of structure and process herein, a peakamplitude PAmp and Jamp of the flow signal derived by filtering from theaccelerometer sensor is used to compute relative changes in the strokevolume. Stroke volume is computed by first applying a blood pressuresignal, or a signal related thereto, to a model of the arterial system.One such model is called a non-linear Windkessel model, which regardsthe blood pressure as analogous to a voltage applied to aseries-parallel network having a series impedance to a output, and aparallel resistor-capacitor combination across the output. These modelcircuit elements are modeled as non-linear to model behavior of thearteries as they expand under blood pressure. The blood flow isanalogous to the voltage across the output of the circuit. Theintegrated output voltage over a period of heart rate S1-to-S1 isrelated to stroke volume for that period and is repeatedly computed.Some other models analogize a reflective electrical transmission line tothe arterial system. Any model appropriate to the purposes at hand isemployed.

Cardiac output CO is defined as the product of stroke volume and heartrate. CO is the volume of blood pumped by the heart per minute. Heartrate is obtained either by counting S1 pulses derived from an axissensor of the accelerometer or counting R pulses using an ECG.

Pre-ejection period PEP is the time interval between onset of ECG QRScomplex and the cardiac ejection. PEP is calculated from the beginningof the ECG QRS complex to the beginning of the first peak in theaccelerometer signal, see FIG. 32. The R-F1 interval approximates thepre-ejection period.

Ventricular contractility VC measures the intrinsic ability of the heartto contract. Contractility can be estimated from the Stroke Volume SV.Increase in Stroke Volume causes an increase in contractility.Contractility VC may alternatively or additionally be measured bytrending the pre-ejection period PEP.

Embodiments of structure and process herein are provided to monitor—bynon-invasive and unobtrusive means—some or all of these vitals andothers. Remarkably, a single, miniature, chest-worn MEMS accelerometeris processed to sense and measure blood flow and other hemodynamicparameters such as stroke volume variations—cardiac output variations,iso-volumic contraction interval; and jointly with a simultaneousECG—contractility and pre-ejection period. The signal corresponding toand related to these parameters is picked up and extracted robustly fromthe accelerometer Y-axis, its axis parallel to or in the plane of thechest and oriented vertically if the patient is standing or seatedvertically, or parallel to a line from head-to-feet (superior-inferior)if the patient is prone or otherwise not standing or seated vertically.

In FIGS. 26 and 30, a system embodiment provides a convenient monitoringdevice and measurement set-up as shown. A miniature (weight—0.08 gram,size—5×5×1.6 mm) triple axis, low-power, analog output MEMSaccelerometer (LIS3L02AL, STMicroelectronics, Geneva, Switzerland) istaped onto the chest and the acceleration signal along theZ-axis—orthogonal to the plane of the chest—corresponding to the heartsounds is captured. Simultaneously, the acceleration signal along theY-axis—in the plane of the chest and oriented verticallyupwards—corresponding to the blood flow and related hemodynamics is alsocaptured by the same MEMS accelerometer. In this way, the accelerationsignal along the in-plane vertical axis (Y-axis) is also captured.

The chest acceleration signals from both axes are, for instance,concurrently AC coupled (high pass rolloff was dropped about 10× in anexample compared to the non-critical three (3) Hz described earlierhereinabove) and separately and in parallel are amplified with a gain of100 and low pass filtered—for anti-aliasing—through a three-stage,5-pole Sallen-and-Key Butterworth filters with a 1 kHz corner frequency.Two commercial quad operational amplifier packages (LT1014CN, LinearTechnology, Milpitas, Calif.) are used for the analog front-end. Theaccelerometer signals are then each sampled at 10,000 (10 K) Samples/secusing a data acquisition card (National Instruments, Austin, Tex.) andcaptured and stored on a computer using MATLAB software (Version 2007b,The Mathworks, Natick, Mass.).

A reference ECG as in FIG. 2 is acquired simultaneously in a threeelectrode (single lead) electrocardiogram configuration for a referencein order to compare with the accelerometer-derived cardiac signal andalso to extract information from the fusion (FIG. 33) of the electricaland mechanical signals (e.g., pre-ejection period PEP, contractilityVC). An additional parallel signal path structure is replicated asdescribed in the previous paragraph and used for the ECG signal.

The extraction of primary heart sounds (S1 and S2—produced by the heartvalve pairs closing at the ends of the diastolic and systolic periodsrespectively of the cardiac cycle) uses a Z-axis sensor of theaccelerometer worn on the chest. In FIG. 31 (and FIG. 4) the primaryheart sounds are robustly detected through post-processing of the Z-axischest acceleration signal as described earlier hereinabove, not onlyduring resting conditions, but also in the presence of stronglyinterfering motion—like walking

The acceleration signal acquired from a subject at rest is digitally lowpass filtered at 50 Hz—using an FIR filter—and decimated by a factor of10. The slow varying respiration baseline wander (e.g., sub-0.5 Hzrespiration and body motion) is removed by smoothing filter andsubtraction to yield a residue signal, and the primary heart sounds (S1and S2) are detected through amplitude and timing based peak detection.

Hemodynamics from Y-axis: As shown in FIG. 27, striking differences wereobserved between the signals picked up by the Y-axis (in the chestplane) and the Z-axis. S1 and S2 align well in shape and timing in bothaxes, but the additional features (lower in frequency with the biggestpeak between S1 and S2—marked at top in black arrows) occur verystrongly along the Y-axis, but are hardly present in the Z-axisdirection, as would be expected of a blood flow-related signal. In thedescription herein, the phrase flow signal or flow-related signal orblood flow signal is applied as a useful identifying label for thischest acceleration component, recognizing that the accelerometer issensing a component of chest acceleration in meters/sec². Thistime-varying acceleration component may also be thought of as abody-reaction acceleration or skin-shear acceleration approximatelyparallel to the superior-inferior body axis in response to blood flowand is related in some way to an overall force of cardiac contraction innewtons and/or to systolic blood pressure in newtons/square meter and/orto blood flow acceleration in milliliters/sec² and/or blood flowvelocity in milliliters/sec. Also, the dynamics of the blood variesspatially and with time at different points in the interior and alongthe blood vessels of the arterial system emanating from the heart.Arterial wall friction and elasticity are involved, and these change ifhardening of the arteries occurs. With these considerations in mind, theaccelerometer signal(s) are processed as described further andpost-processed and interpreted in any appropriate manner by the skilledworker now and in the future to fully realize the benefits of variousembodiments.

For example, the Y-axis signal is filtered or smoothed using a low order(4^(th) order) Savitzky-Golay polynomial filter with a window sizeroughly 200 milliseconds. Different window length and polynomial ordersare feasible. Also, specific polynomials and orders are illustrative andnot limiting because they are related to the signal and the samplingrate. The smoothing filter extracts the slow varying (lower frequency)blood flow signal separated from the residue of the heart sounds (S1 andS2). Using Savitzky-Golay filtering as the smoothing filter is one ofvarious possible ways of extracting the flow signal from the Y-axis.Slow varying respiration and body motion baseline wander (e.g., sub-0.5Hz, below about one-half Hertz) is also removed or separated from theblood flow signal in some embodiments by either cascading or combining ahigh pass filter to attenuate the respiration wander, or using asmoothing filter to isolate the respiration wander and subtracting toyield a residue signal. Further in some embodiments, the respiration isseparated from body motion as taught elsewhere herein. See filteringprocess discussion earlier hereinabove.

FIG. 28 shows a first waveform with the raw Y axis signal and a secondwaveform with the raw Z axis signal. A third waveform traces theextracted flow signal from the S-G filter called a blood flow component,or flow signal, herein. Notice that the 200 msecs window size nicelyencompasses either straight or curved portions within a cycle of theoscillating blood flow signal as the raw Y axis signal stream progressesthrough the filter window. A fourth waveform traces the residue alongthe Y-axis following Savitzky-Golay smoothing (chiefly the S1 and S2sounds). Put another way, FIG. 16 concurrently shows raw Z-axis sensorsignal, raw Y-axis sensor signal, a Savitzky-Golay smoothed signal alongY-axis (interpreted as blood flow), and a signal residue along theY-axis after removal of blood flow component.

In FIG. 29, a first waveform traces the ECG signal. A second waveformshows heart sounds derived from the Z axis accelerometer signal as theresidue of polynomial filtering of FIG. 31 (left side). A third waveformtraces the extracted flow signal from S-G filtering of the accelerometerY-axis sensor signal of FIG. 31 (right side). In this third waveform, aprimary “flow” peak (between S1 and S2) from the accelerometer-basedwaveform is prominent on the third waveform in FIG. 29 and bears somerelationship to a signal from a ballistocardiogram (BCG) while not beingidentical to a BCG. This third waveform is called a flow signal andappears to be related to some extent with the acceleration of surgeblood as it is pushed out of the heart's left ventricle into the aorta.The flow signal obtained from the accelerometer Y-axis has noticeableringy-ness to it, and the additional features of the flow signal of theS-G filtered Y-axis accelerometer sensor can convey some usefulinformation about heart and valve mechanics and possibly other subjects,as described further elsewhere herein. Such information is suitablygenerated and communicated electronically in the wired or wirelessapparatus of FIGS. 24, 25 and 26 operating according to processesdescribed in FIG. 31 and FIG. 36A, and/or in FIG. 36B.

Detection of Isovolumic Contraction Interval IVCI (FIG. 32) is providedin some embodiments by electronically measuring the time intervalbetween the closure of the mitral and tricuspid valve (S1 waveform onthe Z-axis signal) and the start of blood flow (measured by the firstmajor inflection point on the Y-axis flow signal after S1) to output ameasure signal representing the IVCI of the ventricle.

Detection of Pre-ejection Period PEP and Contractility VC involves theQRS waveform of the reference ECG (FIGS. 2, 29, 32) that signifies theinstant of ventricular depolarization. PEP is provided in someembodiments by electronically measuring the time interval between thepeak in the QRS wave and the start of blood flow to output a measuresignal signifying PEP. The electronic measurement jointly processes 1)the flow signal from filtering the accelerometer-derived Y-axis (FIGS.26, 28-31), and 2) the simultaneous ECG as shown in FIG. 2.

In FIG. 31, a process embodiment is represented physically in systemstorage unit or memory of FIG. 25 and executed on the signal processingunit of FIG. 25 or digital signal processor DSP of FIG. 26. In FIG. 31(left side), Z-axis signal processing 110-170 generally is analogous tothe Z-axis processing of FIG. 4, and additionally the peak detection 170of the S1 pulse is followed in FIG. 31 by a time determination 780 of afirst peak location in time P1 of the first heart sound peak S1 in eachheart beat. Counter operation in response to the S1 pulses (and S2pulses or both) electronically generates a heart rate signal as in FIG.36.

In FIG. 31 (right), Y-axis signal processing at step 710 low passfilters (LPFs) the Y-axis input (filter cutoff 50 Hz) and then a step730 S-G polynomial-filters (or otherwise filters effectively) the LPFsignal from step 710 to produce a flow signal output from step 730. Astep 770 performs electronic peak (and trough) detection on the flowsignal to identify a flow peak amplitude PAmp. A succeeding step 785identifies the location F1 of that peak in time, as illustratedgraphically in FIG. 32. Then, in FIG. 31, both the heart rate signaloutput HR and first peak location in time P1 from step 780 are combinedwith and/or compared with first flow signal peak time location F1 fromstep 785 for each heartbeat. The time difference of F1-P1 represents oris proportional to the isovolumic contraction interval IVCI. IVCI inFIGS. 31 and 32 is estimated, for example, as the time differencebetween time P1 of the peak of the S1 waveform (derived from the residuesignal based on the Z-axis sensor) and the time F1 of the first peak ofthe flow signal derived from the Y-axis sensor. (Some embodimentsdetermine heart sound time P1 based on Y-axis processing instead ofZ-axis processing.)

In FIG. 31, Stroke volume SV and relative changes therein are computedat a step 790 proportional to the peak amplitude PAmp of the flow signalderived by filtering from the Y-axis accelerometer sensor. SV in someembodiments is computed proportional to the peak amplitude PAmp of theflow signal multiplied by the ICVI estimate. Cardiac output CO in FIG.31 is electronically generated by multiplying stroke volume SV by heartrate HR derived from Z-axis processing from the step 140 residue fromaccelerometer sensor filtering and the result is supplied to display550.i or 650.i and other devices as discussed in connection with FIG. 25and other system Figures. In FIG. 31, hemodynamic parameters includingall of IVCI, SV, CO, PEP, and VC and others are obtained.

FIG. 32 depicts an ECG, heart sound signal, and flow signal from theaccelerometer 610. The time P1 of occurrence of the S1 sound and thepeak location in time F1 of the flow signal in FIG. 32 are used toprovide a time difference for electronically estimating isovolumiccontraction interval ICVI. Pre-ejection period PEP is electronicallyderived using the time difference between the peak of ECG signal andpeak F1 of the flow signal. A peak amplitude of the flow signal issignified by PAmp in FIG. 32. Peak amplitude PAmp is electronicallyderived as the difference between the highest peak of the flow signalduring a given heartbeat and a baseline average of the flow signal asindicated by a medial dotted line for that heartbeat. A succeeding peakJamp is also depicted. An electronic process embodiment generatessignals representing a value of PEP for every heartbeat and a value ofPAmp for every heartbeat.

FIG. 33 shows superimposed flow signal ensembles for individual heartbeats aligned to the ECG R-waves. The superimposed ensembles showrelative changes (up to 2×) in signal amplitude and timing(jitter/dilation) during recovery of subject from mild exercise. Themultiple ensembles or heart-beat intervals from the same signal acquiredduring exercise recovery are superimposed after aligning to ECG R-wave.They show the amplitude and timing jitter in this Y-axis signal that isinterpreted as blood flow. Amplitude jitter shows that the initial flowis almost twice stronger shortly after exercise and the intensity ofblood flow decreased as the subject recovered over time. Also, thetiming jitter shows that the initial flow peak is closer to the ECGR-wave (shorter pre-ejection period) corresponding to a more contractilestate of the heart. Some embodiments also generate a signal representingthe variance of the jitter in either or both of ICVI and PEP during aperiod of exercise and/or during a period of rest, as well as changes insuch variance value, as an indication of heart function and changestherein.

In some embodiments that measure Pre-ejection period PEP andcontractility VC, one of the ECG electrodes of FIG. 2 is physicallycombined or associated with the accelerometer sensor of FIG. 2, FIG. 26,and/or FIGS. 40A, 40B. Affixing both the electrically-separate ECGsensor and accelerometer as one physical unit simultaneously to thechest affords additional convenience. Another ECG electrode of FIG. 2 insome embodiments has a flexible lead across the chest or along the bodythat physically joins to an operational amplifier chip of FIG. 2 in thesame physical unit as the accelerometer to detect and amplify apotential difference between the two ECG electrodes.

Because of the micro-miniaturization of integrated circuits, thephysical sensor unit is very light in weight and readily taped to thechest. Some further embodiments also include a miniature microphonealong with the accelerometer in the same chest-worn physical unit forobtaining heart sound audio for parallel processing. Various embodimentsrecognize a multitude of concurrent signals that can be obtained by asingle chest accelerometer and provide rich processing to separate themwhile managing to get physiologically relevant information from themultitude of signals. Some of these signals are: Heart-rate,Activity/Motion, Respiration and intrapleural or intra-thoracic pressurechanges, Hemodynamics (timings and amplitudes and changes), Cough,sneeze, snore, speech, breathing sounds, and all other physiologicalprocesses, conditions, and parameters to which the teachings herein lendthemselves. Coughing, sneezing, snoring, speech-related processes, andbreathing sounds are detected in some embodiments by post-processing theaccelerometer for acceleration patterns over both single instances andmultiple instances to distinguish body motions due to coughing,sneezing, snoring, speech-related processes, and breathing sounds fromthose of gait and respiration and other activities.

Detecting and separating coughing, sneezing, snoring, speech-relatedprocesses, and breathing sounds are suitably also or alternativelyprovided by filtering of a microphone input and processing to detect apattern and/or also processing the accelerometer in parallel to otherprocesses described herein and at higher frequencies. In this way,nuanced analysis of cardiovascular, pulmonary, respiratory and otherconditions is conveniently facilitated by the data representing theperspectives that ECG potential difference(s), chest-derived audio, andaccelerometer sensor respectively support.

Some embodiments may be called upon to estimate Pre-ejection period PEPand contractility VC or trends therein, but lack the ECG electrodes ofFIG. 2 and have only the accelerometer sensor. This may be the case inremote monitoring when a person is at a residence and away from theclinic. IVCI can be measured using the accelerometer sensor 610 as theonly data source in some embodiments based on S1-to-F1 interval in FIG.32. Then, recognizing a possibly temporary and not necessarily certainrelationship or correlation of PEP and IVCI, the IVCI is guardedlyeither used as a proxy directly or for trends for Pre-ejection periodPEP (and for use in obtaining contractility VC) where measurementsindicate that it is sufficiently-correlated to PEP or trend therein, orat least sufficient for monitoring variations in PEP and VC to detectexcursion conditions (departures from expected parameter ranges)indicative of advisability of a subsequent clinic visit. Also, previousclinical measurements on the patient with both accelerometer and ECG maybe used to calibrate a time interval adjustment value α and scale valueβ to estimate PEP from IVCI according to Equation (15), wheremeasurements indicate that adjustment α and scale value β arestatistically significant and sufficiently accurate. That adjustment αand scale value β are suitably additionally clinically measured for thepatient as a function of heart rate and any other relevant parameterand, if similarly satisfactory, then further downloaded as a table ofvalues α(HR, etc.) and β(HR, etc.) representing a calibration adjustmentto estimate PEP from IVCI adjusted by such tabulated function α(HR,etc.) and β(HR, etc.) of heart rate HR, etc. That mechanism ofestimating PEP from IVCI involves training a mapping function for eachsubject, recognizing that such approach is likely to be acceptable onlywhen the subject thereafter is to be remote from the clinic and no otheralternative will be immediately available at the remote location.

PEP=β*IVCI+α  (15)

The table of values or parameter table for such functions α(HR, etc.)and β(HR, etc.) is downloaded into flash memory for use by the signalprocessor that processes the accelerometer sensor signal. In that way,PEP estimates and VC estimates are obtained without an ECG andassociated ECG electrodes when the patient is remote from the clinicsubsequently.

FIG. 34 shows the modulation of blood flow parameters—Pre-ejectionperiod PEP (and therefore contractility VC) and flow amplitude PAmp—withexercise recovery. The data is collected for 350 seconds while a subjectrecovers from exercise. A first waveform traces the ECG signal. A secondwaveform traces the extracted flow signal from S-G filtering of theaccelerometer Y-axis sensor signal. A third waveform shows a PEP signalderived from the first and second waveforms. A fourth waveform showspeak flow signal amplitude PAmp of the flow signal of the secondwaveform. As can be seen from the third waveform, the PEP increases (theheart becomes less contractile as subject recovers from exercise).Concurrently, the peak flow signal amplitude PAmp, interpreted as peakblood flow, decreases over time as shown in the fourth waveform. Smallervariations are clarified in the zoomed-in view and likely correspond torespiration modulation of the peak blood flow and contractility. In thisway, the embodiment produces signals as in FIG. 34 for exercise recoverymeasurement of Pre-ejection Period PEP and amplitude of peak blood flowPAmp. Moreover, in some embodiments, respiration signals can beseparated from either PEP or flow peak amplitude PAmp of FIG. 34.

Some medical diagnostic device embodiments have processing embodimentsto detect blood flow using an accelerometer sensor, and hence calculatechanges in various parameters such as Stroke Volume, contractility etc,as described herein. Deriving BCG-like flow data from an accelerometersensor according to embodiments is suitably made part of post-operativerecovery monitoring system embodiments, as well as device embodimentsfor use with an accelerometer for long term, continuous monitoring of apatient's heart. Various embodiments remarkably process input from asingle accelerometer sensor and operate display and therapeutic deviceson the basis of generated signals from the processing thatelectronically represent any, some or all of heart rate, body motion,respiration, blood flow and hemodynamic parameter signals.

In FIGS. 35A and 35B, waveforms are depicted during the Valsalvamaneuver—wherein the subject sits quietly and blows into a mouthpieceunder predetermined back-pressure of the apparatus. By way ofbackground, typical physiological responses expected during the Valsalvamaneuver are that aortic pressure rises from a resting value and thenfalls back to it after a while. Concurrently, heart rate falls belowresting rate and then rises above resting rate. Then breath is releasedafter a predetermined time interval. Aortic pressure again rises fromapproximately the resting value and then falls back to it after a while.Heart rate concurrently falls below the resting rate and then slowlyincreases to the resting rate.

In FIGS. 35A and 35B, actual waveforms during two different instances ofValsalva Release phase are shown. In FIG. 35A, a first waveform tracesthe filtered heart signal residue wherein the heart rate is generallyincreasing, as indicated by decreasing separation between the numerousS1 residue spikes from filtered accelerometer Z-axis as in FIG. 31 (leftside). In FIG. 35B, the first waveform traces the blood flow signal(FIG. 31 right side), and the heart rate is generally increasing also,as indicated by decreasing separation between the numerous flow peaksfrom the flow signal in FIG. 35B. The second waveform of each Figurerepresents peak amplitude PAmp of the first waveform in each sameFigure, which is declining in both instances. The third waveformrepresents declining Stroke Volume, and the fourth waveform representsdeclining Cardiac Output. The waveforms appear to be consistent with thephysiology of the Valsalva maneuver.

In FIGS. 35A/35B, the SV and CO waveforms (3^(th), 4^(th)) were obtainedindirectly, using ModelFlow software, responsive to a separatefinger-mounted sensor using a continuous blood pressure measurementsystem manufactured by Finapres Medical Systems. The system isunderstood to use a non-linear Windkessel model (described elsewhereherein) to model arterial resistance so as to determine blood flow fromcontinuous blood pressure measurements. Notice that theaccelerometer-derived peak amplitude (2^(nd) waveform) in both FIGS. 35Aand 35B tracks the SV and CO.

Accordingly, some embodiments post-process the peak amplitude PAmp(2^(nd) waverform in FIGS. 35A/35B) on Z-axis or other-axis signalamplitude (which correlates well with SV and CO) to provide or derivetime-varying output signals and displays. Such signals and displays anestimation for hemodynamic parameters such as SV and CO and othersderivable directly at FIG. 31 step 790 from the amplitude/power of thecardiac S1 pulse either independently of, or in combination with,information from the blood flow signal. The estimation may differ fromSV and CO themselves by an additive constant and a scale factor, andthis is likely to be acceptable for monitoring applications such asthose that begin with a pre-existing physiological state of a subjectperson and are interested in subsequent variations and/or unusualdepartures. Notice that the SV and CO hemodynamic parameters vary muchmore slowly with time t (e.g., less than 0.2 Hz or less than 0.1 Hz orso) than respiratory variation in the peak amplitude signal PAmp(t), sothat SV and CO are derived from or filtered out of the peak amplitudesignal PAmp(t) in some embodiments and provided for display andrecording. Respiration is separated from the peak amplitude signalPAmp(t) as described in FIG. 22 for instance. Respiration, gait andother body motions are detected and separated from each other based onan accelerometer signal as also taught elsewhere herein and alsoprovided for display and recording.

In FIGS. 36A and 36B, a process embodiment is represented physically insystem storage unit 630 or memory of FIG. 25 and executed on the signalprocessing unit 620 of FIG. 25 or digital signal processor DSP of FIG.26. In FIG. 36A, Z-axis signal processing 110-180 generally is analogousto the Z-axis processing of FIG. 4 and Z-axis processing of FIG. 31(left side), and outputs heart rate at step 180 and need not be furtherdetailed. In FIG. 36B, Y-axis signal processing 910-940, 970, 980independently also derives heart rate by obtaining a residue signal fromthe Y-axis input with the S-G filtered signal subtracted out. Electronicpeak detection of the residue signal at step 940 is followed by peakdetection 970 and counting 980. The heart rate signal output from step980 is either combined with and/or compared with the heart rate outputof FIG. 36A or used instead of and without the heart rate output of FIG.36A, depending on embodiment.

Further in FIG. 36B, the S-G filtered signal from step 930 itself is aringy flow signal of FIGS. 29 and 32 interpreted as blood flow andprovided as an electronic output 950 for display 650.i and optionalstorage 660.i. Also, as described in connection with FIGS. 37 and 38,that flow signal 950 is further processed at a step 960 to recover aForcing Function F(t) as a further electronic output indicative ofcardiac function. In addition, that flow signal is processed at a step965 to estimate one, two, or all of a triplet of 2^(nd) order modelparameters for mass m, dashpot ρ (rho), spring γ (gamma) that aredelivered as still further electronic outputs. FIG. 31 steps 770, 785and 790 are suitably also included in FIG. 36B using the flow signalfrom step 930. Any or all of the outputs can be still furtherpost-processed into electronically-represented interpretations anddisplays in FIG. 25 of the internals of the chest and heart and statesof function.

In FIGS. 37 and 38, post-processing is applied to the damped oscillatoryflow signal at 950 of FIG. 36B (derived from the Y-axis accelerometersensor) as from a 2nd order system model. That 2^(nd) order system modelhas a forcing function F(t) (newtons) and constant coefficientparameters for mass m, dashpot ρ (rho), spring γ (gamma) in its 2ndorder linear differential equation of Equation (16). The variable yrepresents physical displacement of the chest sensor from an average yposition (conceptually measured from some stationary point of referenceon the body, such as the hips, relative to which the chest isdisplaced).

m∂ ² y/∂t ² +ρ∂y/∂t+γy(t)=F(t)  (16)

FIG. 37 models a standing individual with a first triplet of thoseparameters subscripted “1.” FIG. 38 models the individual lying prone,with a second triplet of values for those parameters subscripted “2.” Inboth cases the accelerometer Y-axis sensor is used, where in FIG. 37that sensor is vertical, and in FIG. 38 that sensor is horizontal. Inboth the standing and prone positions that sensor is positioned the sameon the chest, parallel to a superior-inferior axis of symmetry of thebody from head to feet.

Note that the flow signal 950, g(i) derived by step 930 from the Y-axisaccelerometer sensor (e.g. by S-G filtering), can be regarded as aseries of samples g(t) each substantially proportional to the secondderivative ∂²x/∂t² itself in Equation (16). The dashpot parameterρintroduces energy dissipation, and the time constant τ of decay of thedamped oscillatory signal is related to the ratio m/ρ, meaning the massparameter m (kilograms) divided by the dashpot parameter ρ(newtons/(meters/sec)).

τ=m/ρ  (17)

The frequency f_(s) of the damped oscillatory signal is related to(½π)√{square root over ( )}(γ/m), i.e., the square root of the ratio ofthe spring parameter γ (newtons/meter) divided by the mass parameter m(kilograms), and that square-root result divided by 2n.

f _(s)=(½π)√{square root over ( )}(γ/m)  (18)

The post-processing suitably estimates F(t)/m, such as by numericalintegrations directly from the damped oscillatory flow signal waveformfrom the y-axis of the accelerometer, S(t)=∂²y/∂t² using Equation (16)written in the form of Equation (19). The numerical integration beginsas each spindle-shaped accelerometer Y-axis waveform commences in FIG.29 (3rd waveform) for a given heartbeat, and assumes that any constantsof integration are zero (i.e., zero position, zero velocity. Forapplications based on the shape or morphology of the forcing functionF(t), the mass m is merely a constant of proportionality that does notaffect the shape. If mass is important to the application, the mass istaken as that of the head and torso such as some fraction (e.g., 0.6) ofthe body mass in kilograms. The time constant τ is numerically estimatedas the length of time from the peak of the spindle-shaped accelerationwaveform to the time when the waveform is about ⅔ dissipated (i.e.,reduced on later end of the spindle to 1/e of its earlier peakamplitude, where e is base of natural logarithms 2.71828 . . . ). InFIG. 29, the time constant is about a quarter of a second. The frequencyf_(s) is numerically estimated as the number of cycles in some portionof the spindle-shaped acceleration waveform divided by the time inseconds occupied by that portion. In FIG. 29, the frequency f_(s) isabout 8 Hertz.

F(t)/m=S(t)+(1/τ)∫₀ ^(t) S(t)dt+(2πf _(s))²∫₀ ^(t)∫₀ ^(t) S(t)dt  (19)

Alternatively, the post-processing uses any applicable statisticaltime-series analysis package or procedure to recover best statisticalestimates for the forcing function and the 2^(nd) order constantcoefficient parameters.

The forcing function F_(Y) (t) component parallel to the Y-axis sensormay arise from a mixture of 1) physical acceleration of the heart itselfupon ventricular contraction and 2) the acceleration of blood surginginto the aorta when the blood is expelled from the left ventricle. Theparameter γ for spring-constant and parameter ρ for dashpot seem torelate to some gross average of mechanical properties of the interiorsof chest and abdomen. The mass parameter m probably is related orproportional to the mass of the torso and perhaps the head, but probablynot to the mass of the legs because the legs are probably notaccelerated in the Y-axis direction. The observed S1-S1 waveform alsohas a rising amplitude of oscillation immediately preceding the dampedoscillation, see FIG. 32. The latter behavior can be due to entry ofblood from the venae cavae into the right atrium, and the rightventricular contraction into the pulmonary artery. Accordingly,embodiments for extraction and analysis of forcing function F_(Y)(t) canprovide useful and more nearly comprehensive information on cardiac andpulmonary function as well as hemodynamic information.

Some embodiments are contemplated that monitor accelerometer X-axissensor information as well as the Y-axis and Z-axis. By X-axis sensor ismeant a sensor oriented to sense acceleration laterally across thechest. A transverse displacement variable x for purposes of Equation(20-X) represents side-to-side physical displacement of the chest sensorfrom an average x position (or conceptually also from a point ofreference such as the center of mass of the heart relative to which thechest is displaced.) In such an embodiment, signal from the X-axissensor is filtered in parallel with the filtering of the Y-axis signal,and in a manner for the X-axis analogous to the filtering describedhereinabove for the Y-axis signal. Because of the assymetrical locationand slantwise inclination and of the heart in the chest, the filteredsignal from the X-axis sensor provides further information about alateral (side-to-side) component F_(X)(t) of the forcing function F(t)considered as a vector. Taken together, these two forcing functioncomponents F_(Y)(t) and F_(X)(t) can provide further useful informationon cardiac function, pulmonary function, properties of the pleura,pleural cavity, and pericardium, as well as hemodynamics informationrelating to the aorta, venae cavae, and pulmonary arteries and pulmonaryveins by any suitable process now known or hereafter devised. Theparameter triplets are respectively subscripted “1Y” and “1X” todesignate a standing position (“1”) and the Y-axis or X-axis sensorinvolved. If the prone position is involved then the subscript “1” ischanged to “2.”

m _(1Y)∂² y/∂t ²+ρ_(1Y) ∂y/∂t+γ _(1Y) y(t)=F _(1Y)(t)  (20-Y)

m _(1X)∂² x/∂t ²+ρ_(1X) ∂x/∂t+γ _(1X) x(t)=F _(1X)(t).  (20-X)

In FIG. 39, an embodiment is improved as in the other Figures herein andused as one or more replicas as discussed in connection with FIG. 25.FIG. 39 illustrates inventive integrated circuit chips including chips1100, 1200, 1300, 1400, 1500, and GPS 1190 (1495) for use in any one,some or all of the blocks of the communications system 600 of FIG. 25.The skilled worker uses and adapts the integrated circuits to theparticular parts of the communications system 600 as appropriate to thefunctions intended. It is contemplated that the skilled worker uses eachof the integrated circuits shown in FIG. 39, or such selection from thecomplement of blocks therein provided into appropriate other integratedcircuit chips, or provided into one single integrated circuit chip, in amanner optimally combined or partitioned between the chips, to theextent needed by any of the applications supported such as voice WLANgateway, cellular telephones, televisions, internet audio/video devices,routers, pagers, personal digital assistants (PDA), microcontrollerscoupled to controlled mechanisms for fixed, mobile, personal, roboticand/or automotive use, combinations thereof, and other applicationproducts now known or hereafter devised for increased, partitioned orselectively determinable advantages.

In FIG. 39, an integrated circuit 1100 includes a digital baseband (DBB)block that has a RISC processor 1105 (such as MIPS core(s), ARM core(s),or other suitable processor) and a digital signal processor 1110 such asfrom the TMS320C55x™ DSP generation from Texas Instruments Incorporatedor other digital signal processor (or DSP core) 1110, communicationssoftware and security software for any such processor or core, securityaccelerators 1140, and a memory controller. Security accelerators 1140provide additional computing power such as for hashing and encryptionthat are accessible, for instance, when the integrated circuit 1100 isoperated in a security level enabling the security accelerators block1140 and affording types of access to the security acceleratorsdepending on the security level and/or security mode. The memorycontroller interfaces the RISC core 1105 and the DSP core 1110 to Flashmemory 1025 and SDRAM 1024 (synchronous dynamic random access memory).On chip RAM 1120 and on-chip ROM 1130 also are accessible to theprocessors 1105 and 1110 for providing sequences of softwareinstructions and data thereto. A security logic circuit 1038 of FIGS. 16and 17 has a secure state machine (SSM) to provide hardware monitoringof any tampering with security features. A Secure Demand Paging (SDP)circuit 1040 is provided for effectively-extended secure memory.

Digital circuitry 1150 on integrated circuit 1100 supports and provideswireless modem interfaces for any one or more of GSM, GPRS, EDGE, UMTS,and OFDMA/MIMO (Global System for Mobile communications, General PacketRadio Service, Enhanced Data Rates for Global Evolution, UniversalMobile Telecommunications System, Orthogonal Frequency Division MultipleAccess and Multiple Input Multiple Output Antennas) wireless, with orwithout high speed digital data service, via an analog baseband chip1200 and GSM/CDMA transmit/receive chip 1300. Digital circuitry 1150includes a ciphering processor CRYPT for GSM ciphering and/or otherencryption/decryption purposes. Blocks TPU (Time Processing Unitreal-time sequencer), TSP (Time Serial Port), GEA (GPRS EncryptionAlgorithm block for ciphering at LLC logical link layer), RIF (RadioInterface), and SPI (Serial Port Interface) are included in digitalcircuitry 1150.

Digital circuitry 1160 provides codec for CDMA (Code Division MultipleAccess), CDMA2000, and/or WCDMA (wideband CDMA or UMTS) wirelesssuitably with HSDPA/HSUPA (High Speed Downlink Packet Access, High SpeedUplink Packet Access) (or 1xEV-DV, 1xEV-DO or 3xEV-DV) data feature viathe analog baseband chip 1200 and RF GSM/CDMA chip 1300. Digitalcircuitry 1160 includes blocks MRC (maximal ratio combiner for multipathsymbol combining), ENC (encryption/decryption), RX (downlink receivechannel decoding, de-interleaving, viterbi decoding and turbo decoding)and TX (uplink transmit convolutional encoding, turbo encoding,interleaving and channelizing.). Blocks for uplink and downlinkprocesses of WCDMA are provided.

Audio/voice block 1170 supports audio and voice functions andinterfacing. Speech/voice codec(s) and speech recognition are suitablyprovided in memory space in audio/voice block 1170 for processing byprocessor(s) 1110. An applications interface block 1180 couples thedigital baseband chip 1100 to an applications processor 1400. Also, aserial interface in block 1180 interfaces from parallel digital buses onchip 1100 to USB (Universal Serial Bus) of PC (personal computer) 2070.The serial interface includes UARTs (universal asynchronousreceiver/transmitter circuit) for performing the conversion of databetween parallel and serial lines. A power resets and control modulePRCM 1185 provides power management circuitry for chip 1100. Chip 1100is coupled to location-determining circuitry 1190 satellite positioningsuch as GPS (Global Positioning System) and/or to a network-basedpositioning (triangulation) system, to an accelerometer, to a tiltsensor, and/or other peripherals to support positioning, position-basedapplications, user real-time kinematics-based applications, and othersuch applications. Chip 1100 is also coupled to a USIM (UMTS SubscriberIdentity Module) 1195 or other SIM for user insertion of an identifyingplastic card, or other storage element, or for sensing biometricinformation to identify the user and activate features.

In FIG. 39, a mixed-signal integrated circuit 1200 includes an analogbaseband (ABB) block 1210 for GSM/GPRS/EDGE/UMTS/HSDPA/HSUPA whichincludes SPI (Serial Port Interface),digital-to-analog/analog-to-digital conversion DAC/ADC block, and RF(radio frequency) Control pertaining to GSM/GPRS/EDGE/UMTS/HSDPA/HSUPAand coupled to RF (GSM etc.) chip 1300. Block 1210 suitably provides ananalogous ABB for CDMA wireless and any associated 1xEV-DV, 1xEV-DO or3xEV-DV data and/or voice with its respective SPI (Serial PortInterface), digital-to-analog conversion DAC/ADC block, and RF Controlpertaining to CDMA and coupled to RF (CDMA) chip 1300.

An audio block 1220 has audio I/O (input/output) circuits to a speaker1222, a microphone 1224, and headphones (not shown). Audio block 1220has an analog-to-digital converter (ADC) coupled to an audio/voice codec1170 and a stereo DAC (digital to analog converter) for a signal path tothe baseband block 1210 and with suitable encryption/decryption. Acontrol interface 1230 has a primary host interface (I/F) and asecondary host interface to DBB-related integrated circuit 1100 of FIG.39 for the respective GSM and CDMA paths. The integrated circuit 1200 isalso interfaced to an I2C port of applications processor chip 1400 ofFIG. 39. Control interface 1230 is also coupled via circuitry tointerfaces in circuits 1250 and the baseband 1210. A power conversionblock 1240 includes buck voltage conversion circuitry for DC-to-DCconversion, and low-dropout (LDO) voltage regulators for powermanagement/sleep mode of respective parts of the chip regulated by theLDOs. Power conversion block 1240 provides information to and isresponsive to a power control state machine between the power conversionblock 1240 and circuits 1250. Power management circuitry PRCM 1185(1470) is coupled with and controls power conversion block 1240 andinterfaces to GPS 1190 (1495) and to power save mode controller 2130(2290) in systems of FIGS. 1-39 and as described elsewhere herein.Circuits 1250 provide oscillator circuitry for clocking chip 1200. Theoscillators have frequencies determined by one or more crystals 1290.Circuits 1250 include a RTC real time clock (time/date functions),general purpose I/O, a vibrator drive (supplement to cell phone ringingfeatures), and a USB On-The-Go (OTG) transceiver. A touch screeninterface 1260 is coupled to a touch screen XY 1266 off-chip. Batteriessuch as a lithium-ion battery 1280 and backup battery and rechargerprovide power to the system and battery data to circuit 1250 on suitablyprovided separate lines from the battery pack. When needed, the battery1280 also receives charging current from a Charge Controller in analogcircuit 1250 which includes MADC (Monitoring ADC and analog inputmultiplexer such as for on-chip charging voltage and current, andbattery voltage lines, and off-chip battery voltage, current,temperature) under control of the power control state machine. Batterymonitoring is provided by either or both of 1-Wire and/or an interfacecalled HDQ.

In FIG. 39 an RF integrated circuit 1300 includes aGSM/GPRS/EDGE/UMTS/CDMA RF transmitter block 1310 supported byoscillator circuitry with crystal(s) 1290. Transmitter block 1310 is fedby basebands block 1210 of chip 1200. Transmitter block 1310 drives adual band RF power amplifier (PA) 1330. On-chip voltage regulatorsmaintain appropriate voltage under conditions of varying power usage.Off-chip switchplexer 1350 couples wireless antenna and switch circuitryto both the transmit portion 1310, 1330 and the receive portion nextdescribed. Switchplexer 1350 is coupled via band-pass filters 1360 toreceiving LNAs (low noise amplifiers) for 850/900 MHz, 1800 MHz, 1900MHz and other frequency bands as appropriate. Depending on the band inuse, the output of LNAs couples to GSM/GPRS/EDGE/UMTS/CDMA demodulator1370 to produce the I/Q or other outputs thereof (in-phase, quadrature)to the GSM/GPRS/EDGE/UMTS/CDMA basebands block 1210.

Further in FIG. 39, an integrated circuit chip or core 1400 is providedfor applications processing and more off-chip peripherals. Chip (orcore) 1400 has interface circuit 1410 including a high-speed WLAN802.11a/b/g interface coupled to a WLAN chip 1500. Further provided onchip 1400 is an applications processing section 1420 which includes aRISC processor 1422 (such as MIPS core(s), ARM core(s), or othersuitable processor), a digital signal processor (DSP) 1424 such as fromthe TMS320C55x™ DSP generation and/or the TMS320C6x™ DSP generation fromTexas Instruments Incorporated or other digital signal processor(s), anda shared memory controller MEM CTRL 1426 with DMA (direct memoryaccess), and a 2D (two-dimensional display) graphic accelerator.Speech/voice codec/speech recognition functionality is suitablyprocessed in chip 1400, in chip 1100, or both chips 1400 and 1100.

The RISC processor 1422 and the DSP 1424 in section 1420 have access viaan on-chip extended memory interface (EMIF/CF) to off-chip memoryresources 1435 including as appropriate, mobile DDR (double data rate)DRAM, and flash memory of any of NAND Flash, NOR Flash, and CompactFlash. On chip 1400, a shared memory controller 1426 in circuitry 1420interfaces the RISC processor 1420 and the DSP 1424 via an on-chip busto on-chip memory 1440 with RAM and ROM. A 2D graphic accelerator iscoupled to frame buffer internal SRAM (static random access memory) inblock 1440. A security block 1450 includes an SSM analogous to SSM 1038of FIG. 1, and includes secure hardware accelerators having securityfeatures and provided for secure demand paging 1040 and for acceleratingencryption and decryption. A random number generator RNG is provided insecurity block 1450.

On-chip peripherals and additional interfaces 1410 include UART datainterface and MCSI (Multi-Channel Serial Interface) voice and datawireless interface for an off-chip IEEE 802.15 (Bluetooth and low andhigh rate piconet, Zigbee, and personal network communications) wirelesscircuit 1430. The Bluetooth or Zigbee wireless interface is useful forreceiving from and controlling the accelerometer sensor and itsassociated analog circuitry and digital to analog-to-digital converterADC in FIGS. 1 and 26, among other Figures. In arrangements includingECG electrodes and/or a chest microphone, the analog circuitry at thetaped-on sensor unit also includes couplings from such pickup elementsto the Bluetooth or Zigbee short distance transceiver from the chestsensor (e.g. FIGS. 40A/40B communicating with a counterpart shortdistance transceiver at the interface 1410.

Debug messaging and serial interfacing are also available through theUART. A JTAG emulation interface couples to an off-chip emulatorDebugger for test and debug. GPS 1190 (1495) is scannable by thedebugger, see FIG. 2. Further in peripherals 1410 are an I2C interfaceto analog baseband ABB chip 1200, and an interface to applicationsinterface 1180 of integrated circuit chip 1100 having digital basebandDBB.

Interface 1410 includes a MCSI voice interface, a UART interface forcontrols and data to position unit GPS 1495 and otherwise, and amulti-channel buffered serial port (McBSP) for data. Timers, interruptcontroller, and RTC (real time clock) circuitry are provided in chip1400. Further in peripherals 1410 are a MicroWire (u-wire 4 channelserial port) and multi-channel buffered serial port (McBSP) to Audiocodec, a touch-screen controller (or coupling to 1260), and audioamplifier 1480 to stereo speakers.

External audio content and touch screen (in/out) 1260, 1266 and LCD(liquid crystal display), organic semiconductor display, and DLP™digital light processor display from Texas Instruments Incorporated, aresuitably provided in various embodiments and coupled to interface 1410.In vehicular use, such as at unit 690 of FIG. 25, the display issuitably any of these types provided in the vehicle, and sound isprovided through loudspeakers, headphones or other audio transducersprovided in the vehicle. In some vehicles a transparent organicsemiconductor display 2095 of FIG. 16 is provided on one or more windowsof a vehicle and wirelessly or wireline-coupled to the video feed. Mapsand visual position-based interactive imaging and user kinematicsapplications are provided using double-integrated accelerometer outputas discussed elsewhere herein. Also GPS 1190 (1495) and processor 1105,1110 (1422, 1424) support fixed, portable, mobile, vehicular and otherplatforms.

Interface 1410 additionally has an on-chip USB OTG interface thatcouples to off-chip Host and Client devices. These USB communicationsare suitably directed outside handset 2010 such as to PC 2070 (personalcomputer) and/or from PC 2070 to update the handset 2010 or to a camera1490.

An on-chip UART/IrDA (infrared data) interface in interfaces 1410couples to off-chip GPS (global positioning system of block 1495cooperating with or instead of GPS 1190) and Fast IrDA infrared wirelesscommunications device. An interface provides EMT9 and Camera interfacingto one or more off-chip still cameras or video cameras 1490, and/or to aCMOS sensor of radiant energy. Such cameras and other apparatus all haveadditional processing performed with greater speed and efficiency in thecameras and apparatus and in mobile devices coupled to them withimprovements as described herein. Further in FIG. 39, an on-chip LCDcontroller or DLP™ controller and associated PWL (Pulse-Width Light)block in interfaces 1410 are coupled to a color LCD display or DLP™display and its LCD light controller off-chip and/or DLP™ digital lightprocessor display.

Further, on-chip interfaces 1410 are respectively provided for off-chipkeypad and GPIO (general purpose input/output). On-chip LPG (LED PulseGenerator) and PWT (Pulse-Width Tone) interfaces are respectivelyprovided for off-chip LED and buzzer peripherals. On-chip MMC/SDmultimedia and flash interfaces are provided for off-chip MMC Flashcard, SD flash card and SDIO peripherals. On chip 1400, a power, resets,and control module PRCM 1470 supervises and controls power consumingblocks and sequences them, and coordinates with PRCM 1185 on chip 1100and with Power Save Mode Controller 2130 (2290) in GPS 1495 as describedelsewhere herein.

In FIG. 39, a WLAN integrated circuit 1500 includes MAC (media accesscontroller) 1510, PHY (physical layer) 1520 and AFE (analog front end)1530 for use in various WLAN and UMA (Unlicensed Mobile Access) modemapplications. In some embodiments, GPS 1495 operates in closecoordination with any one, some, or all of WLAN, WiMax, DVB, or othernetwork, to provide positioning, position-based, and user real-timekinematics applications. Still other additional wireless interfaces suchas for wideband wireless such as IEEE 802.16 WiMAX mesh networking andother standards are suitably provided and coupled to the applicationsprocessor integrated circuit 1400 and other processors in the system.WiMax has MAC and PHY processes and the illustration of blocks 1510 and1520 for WLAN indicates the relative positions of the MAC and PHY blocksfor WiMax.

In FIG. 39, a further digital video integrated circuit 1610 is coupledwith a television antenna 1615 (and/or coupling circuitry to shareantenna 1015 and/or 1545 and/or 2105) to provide television antennatuning, antenna selection, filtering, RF input stage for recoveringvideo/audio/controls from television transmitter (e.g., DVB station 2020of FIG. 16). Digital video integrated circuit 1610 in some embodimentshas an integrated analog-to-digital converter ADC on-chip, and in someother embodiments feeds analog to ABB chip 1200 for conversion by an ADCon ABB chip 1200. The ADC supplies a digital output 1619 to interfaces1410 of applications processor chip 1400 either directly from chip 1610or indirectly from chip 1610 via the ADC on ABB chip 1200. Controls forchip 1610 are provided on lines 1625 from interfaces 1410. Applicationsprocessor chip 1400 includes a digital video block 1620 coupled tointerface 1410 and having a configurable adjustable shared-memorytelecommunications signal processing chain such as Doppler/MPE-FEC. Aprocessor on chip 1400 such as RISC processor 1422 and/or DSP 1424configures, supervises and controls the operations of the digital videoblock 1620.

In combination with the GPS circuit 1190 and/or 1495, and video display1266 or LCD, the RISC processor 1105/1422 and/or DSP 1110 (1424) supportlocation-based embodiments and services of various types, such asroadmaps and directions thereon to a destination, pictorials of nearbycommercial establishments, offices, and residences of friends, variousfamily supervision applications, position sending to friends or toemergency E911 service, and other location based services now known oryet to be devised.

Digital signal processor cores suitable for some embodiments in the IVAblock and video codec block may include a Texas Instruments TMS32055x™series digital signal processor with low power dissipation, and/orTMS320C6000 series and/or TMS320C64x™ series VLIW digital signalprocessor, and have the circuitry and processes of the FIGS. 1-39coupled with them as taught herein. A camera CAM provides video pickupfor a cell phone or other device to send over the internet to anothercell phone, personal digital assistant/personal entertainment unit,gateway and/or set top box STB with television TV.

FIGS. 40A and 40B are respective broadside and cross-sectional views ofan accelerometer sensor 210 and transmitter, transceiver, or transponderchip 212 firmly mounted on a thin, resilient plastic support plate 214that can be firmly affixed by an adhesive tape 216 to the chest. Theelectronics is conveniently light-weight and small and may bequarter-sized, dime-sized or even smaller in size. In FIG. 40A, a dottedoutline shows a round smoothed or flanged periphery of plastic support214 shaped for comfort on the chest.

In FIG. 40B, the Z-axis of accelerometer sensor 210 is perpendicular tothe plane defined by plastic support 214 (and to the plane defined bychip 212). The Y-axis and X-axis of sensitivity to acceleration of theaccelerometer sensor 210 are perpendicular to each other, with eachparallel to the plane defined by a broadside of a package enclosing theaccelerometer and likewise parallel to a plane defined by plasticsupport 214. Adhesive tape 216 adheres to the outward broad side ofplastic support 214, thereby holding plastic support 214 firmly againstthe chest when applied thereto. Adhesive tape 216 has an inner edge 217defining an approximately square aperture in FIG. 40A that admits theoutward-placed transponder chip 212 and accelerometer sensor 210.

An ECG sensor of FIG. 2 and/or a small microphone may also be mounted onplastic support 214 to monitor chest potential and/or chest sounds. Thechest-adjacent side of plastic support 214 may also be provided with ECGelectrode paste for ECG connectivity with the chest.

In some embodiments, chip 212 harvests power from an interrogationsignal from the circuitry of FIG. 39, and in other embodiments a smallbattery is also provided on plastic support 214 and electricallyconnected to supply a low power to chip 212. Accelerometer sensor 210 iselectrically coupled to transponder chip 212 along with any ECGelectrode and microphone elements for wireless communication to thesystem of FIG. 39. In various embodiments, none, one, some or all of theblocks of FIG. 1 are provided as part of transponder chip 212. Chip 212in some embodiments includes a very low power processor such as anMSP430™ processor from Texas Instruments Incorporated or other suchprocessor along with the short distance wireless transmitter. Chip 212can have an antenna such as a spiral antenna fabricated as part of thechip 212, or in some other embodiments an antenna is suitably providedas part of plastic support 214 and electrically connected to chip 212.Optionally, a plastic cap or header physically encloses and protects thechips over the support 214. Also, in some embodiments, a wirelineinterface is also provided in chip 212, and the support 214 physicallyhas a miniature wireline female connector attached thereto andelectrically connected to the wireline interface in chip 212, such asfor USB (Universal Serial Bus). In that way, a clinician may connect alightweight male connector from a monitoring processor and display unitto the miniature wireline female connector and bypass the short distancewireless function of chip 212 at will. In still other embodiments, theaccelerometer 210 and transponder 212 are mounted in a pacemaker that iseither implanted in the patient or affixed to the chest.

In FIGS. 40A and 40B, the orientation of the X, Y, and Z axes of theaccelerometer sensor on the chest may vary depending on actual placementand actual physical manufacture. Actual orientation of the accelerometersensor on the chest may vary because of convenience for categories ofpatients or particular patients or simply due to inadvertentmis-orientation of the sensor. However, physical orientation of themultiple axis accelerometer merely distributes the overall physicalacceleration vector a to be sensed to the different sensor axes of theaccelerometer according to their vector components in the various axisdirections.

Accordingly, some embodiments as in FIG. 41 include an electronicprocessing module 990.i for virtual re-orientation or optimization ofthe accelerometer sensor signals by applying a rotation of axes thatintroduces a multiplication by a rotation matrix to the signals. Suchrotation of axes 990.i combines the signals for X, Y, Z accelerometeraxes, such as shown for those axes in FIG. 42, according to linearcombinations of the signals as if the accelerometer axes were rotated,as shown by Equation (21).

Let an angle θ represent an angle by which the accelerometer Y-axissensor is to be virtually rotated from its affixed position on the chestto align with the foot-to-head direction on the body or for whateverpurpose. Let an angle φ represent an angle by which the accelerometerZ-axis sensor is to be virtually rotated from its affixed positionapproximately perpendicular to the chest toward that foot-to-headdirection on the body. Let a vector V represent the Z-axis signal, theY-axis signal and the X-axis signal. Vector V of these signals is matrixmultiplied electronically in FIG. 41 according to the rotation productR*V, using rotation matrix R expressed by Equation (21).

$\begin{matrix}{R = \begin{bmatrix}{\cos \; \phi} & \left( {\sin \; \phi \; \sin \; \theta} \right) & \left( {\sin \; \phi \; \cos \; \theta} \right) \\0 & {\cos \; \theta} & {{- \sin}\; \theta} \\{{- \sin}\; \phi} & \left( {\cos \; \phi \; \sin \; \theta} \right) & \left( {\cos \; {\phi cos}\; \theta} \right)\end{bmatrix}} & (21)\end{matrix}$

In various embodiments, the axis rotations are suitably customized bythe processing for the type of signal output (e.g., blood flow, heartsounds) which is to be maximized for a given purpose. The angles θ and φare each varied by a given feedback control circuit 995.i to maximizethe desired type of signal output to which that feedback control circuitis applied.

In FIG. 41, for instance, some embodiments execute a feedback control995.1 to thus maximize the heart sound signal for the heart monitoringpath, and the axes-rotation parameters for the rotation process 990.1for a feedback loop 990.1, 130, 140, . . . , 995.1 established either asa configuration routine before run-time, or dynamically at run-time. Thefeedback loop rotates the axes to deliver a linear combination of Z-axisand Y-axis (and X-axis can also be useful) as an input in place of theraw Z-axis signal in FIGS. 31 and 36A to the Savitzky-Golay polynomialfilter 130 for the Z-axis to maximize the amplitude of the S1 peaks atthe output of the Folded Correlation, for heart sound and heart ratemonitoring purposes.

Analogously, in FIG. 41, some embodiments additionally or alternativelyexecute a feedback loop 990.2, 930, . . . , 995.2 by independentlyrotating axes to deliver a linear combination of X-axis and Y-axis (andZ-axis can also be useful) as an input in place of the raw Y-axis signalin FIGS. 31 and 36B to the Savitzky-Golay polynomial filter 930 for theY-axis, to maximize the blood flow signal peak amplitude PAmp of FIGS.29 and 32. To save some processing, some embodiments can perform theblood flow axes-rotation on the X and Y axes only (φ=0 in Equation (21))for the blood flow signal, see Equation (22).

$\begin{matrix}{{R\left( {\phi = 0} \right)} = \begin{bmatrix}1 & 0 & 0 \\0 & {\cos \; \theta} & {{- \sin}\; \theta} \\0 & {\sin \; \theta} & {\cos \; \theta}\end{bmatrix}} & (22)\end{matrix}$

Note that the rotation matrix R of Equation (21) is the product of atilt matrix M of Equation (23) with the XY rotation matrix of Equation(22):

$\begin{matrix}{M = \begin{bmatrix}{\cos \; \phi} & 0 & {\sin \; \phi} \\0 & 1 & 0 \\{{- \sin}\; \phi} & 0 & {\cos \; \phi}\end{bmatrix}} & (23)\end{matrix}$

In another way to save some processing, some embodiments can use onerotation 990 and one feedback control 995 operating in response tosignals jointly, like heart sounds amplitude and/or blood flow signalamplitude. Various modes of operation and configuration can be activatedor disabled by means of one or more control registers with bits or bitfields for the various operations and configurations. A manual mode, ifactivated, can override the feedback controls and let a clinicianmanually optimize the virtual rotations while examining signals likethose of FIG. 29 on the computer display of FIG. 25.

Some embodiments also include an electronic compass physically includedinto the assembly of FIGS. 40A, 40B for supporting location-basedservices by the sensor assembly. An e-compass and signals therefrom areprovided, calibrated and processed using the teachings of US patentapplication “Processes for More Accurately Calibrating E-Compass forTilt Error, Circuits, and Systems” Ser. No. 12/398,696 (TI-65997) filedMar. 5, 2009, and which is incorporated herein by reference in itsentirety.

FIG. 42 shows four concurrent waveforms including reference ECG,acceleration along the dorso-ventral axis (Z-axis), acceleration alongthe superior-inferior axis (Y-axis) and acceleration along thedextro-sinistral axis (X-axis). Notice the relatively prominent heartpeaks in the Z-axis and X-axis waveforms and the relatively prominentspindle-shaped oscillatory blood flow component in the Y-axis waveform.The latter three acceleration signals for X, Y, Z accelerometer axes aresuitably applied in the circuit FIG. 41 and circuits and processes ofany other Figures that can benefit from use of signals from two or threeof the accelerometer axes. Some further embodiments provide circuitryand/or firmware that fuses hemodynamic and acoustic signatures frommultiple-axis signals and/or inter-axis cross-talk using approaches likeblind source separation, principal component analysis PCA and/orindependent component analysis ICA.

Various embodiments as described herein are manufactured in a processthat prepares a particular design and printed wiring board (PWB) of thesystem unit and has an applications processor coupled to a modem,together with one or more peripherals coupled to the processor and auser interface coupled to the processor or not, as the case may be. Astorage, such as SDRAM and Flash memory is coupled to the system (e.g.,FIG. 39) and has tables, configuration and parameters and an operatingsystem OS, protected applications (PPAs and PAs), and other supervisorysoftware. System testing tests operations of the integrated circuit(s)and system in actual application for efficiency and satisfactoryoperation of fixed or mobile video display for continuity of datatransfer and content, display and other user interface operation andother such operation that is apparent to the human user and can beevaluated by system use. If further increased efficiency is called for,parameter(s) are reconfigured for further testing. Adjusted parameter(s)are loaded into the Flash memory or otherwise, components are assembledon PWB to produce resulting system units.

The electronic monitoring devices and processing described herein issuitably supported by any one or more of RISC (reduced instruction setcomputing), CISC (complex instruction set computing), DSP (digitalsignal processors), microcontrollers, PC (personal computer) mainmicroprocessors, math coprocessors, VLIW (very long instruction word),SIMD (single instruction multiple data) and MIMD (multiple instructionmultiple data) processors and coprocessors as cores or standaloneintegrated circuits, and in other integrated circuits and arrays. Othertypes of integrated circuits are applied, such as ASICs (applicationspecific integrated circuits) and gate arrays and all circuits to whichthe advantages of the improvements described herein commend their use.

In addition to inventive structures, devices, apparatus and systems,processes are represented and described using any and all of the blockdiagrams, logic diagrams, and flow diagrams herein. Block diagram blocksare used to represent both structures as understood by those of ordinaryskill in the art as well as process steps and portions of process flows.Similarly, logic elements in the diagrams represent both electronicstructures and process steps and portions of process flows. Flow diagramsymbols herein represent process steps and portions of process flows insoftware and hardware embodiments as well as portions of structure invarious embodiments of the invention.

ASPECTS See Notes Paragraph at End of this Aspects Section

Notes about Aspects above: Aspects are paragraphs which might be offeredas claims in patent prosecution. The above dependently-written Aspectshave leading digits and internal dependency designations to indicate theclaims or aspects to which they pertain. Aspects having no internaldependency designations have leading digits and alphanumerics toindicate the position in the ordering of claims at which they might besituated if offered as claims in prosecution.

Processing circuitry comprehends digital, analog and mixed signal(digital/analog) integrated circuits, ASIC circuits, PALs, PLAs,decoders, memories, and programmable and nonprogrammable processors,microcontrollers and other circuitry. Internal and external couplingsand connections can be ohmic, capacitive, inductive, photonic, anddirect or indirect via intervening circuits or otherwise as desirable.Process diagrams herein are representative of flow diagrams foroperations of any embodiments whether of hardware, software, orfirmware, and processes of manufacture thereof. Flow diagrams and blockdiagrams are each interpretable as representing structure and/orprocess. While this invention has been described with reference toillustrative embodiments, this description is not to be construed in alimiting sense. Various modifications and combinations of theillustrative embodiments, as well as other embodiments of the inventionmay be made. The terms including, includes, having, has, with, orvariants thereof are used in the detailed description and/or the claimsto denote non-exhaustive inclusion in a manner similar to the termcomprising. The appended claims and their equivalents should beinterpreted to cover any such embodiments, modifications, andembodiments as fall within the scope of the invention.

1. An electronic monitoring device comprising: an electronic processorhaving at least one signal input for body monitoring; and a memoryholding instructions for said electronic processor coupled to saidelectronic processor so that said electronic processor is operable toisolate a cardiac signal including cardiac pulses combined with othercardiac signal variations, and said electronic processor furtheroperable to execute a filter that separates a varying blood flow signalfrom the cardiac pulses and to output information based on at least thevarying blood flow signal.
 2. The electronic monitoring device claimedin claim 1 wherein said electronic processor is operable to alsoseparate cardiac pulses from the cardiac signal and produce anelectronic representation of a time interval between a selected cardiacpulse and an event on the varying blood flow signal.
 3. The electronicmonitoring device claimed in claim 2 wherein the selected cardiac pulseis an S1 pulse.
 4. The electronic monitoring device claimed in claim 2wherein the event is a peak in the varying blood flow signal after theselected cardiac pulse.
 5. The electronic monitoring device claimed inclaim 2 wherein said electronic processor is operable to produce ameasurement of jitter across heart beats of the time interval.
 6. Theelectronic monitoring device claimed in claim 1 wherein said electronicprocessor is operable to output an electronic representation ofiso-volumic contraction interval using the varying blood flow signal. 7.The electronic monitoring device claimed in claim 6 wherein saidelectronic processor is operable to produce an electronic representationof contractility using the iso-volumic contraction interval.
 8. Theelectronic monitoring device claimed in claim 1 wherein the filter forsaid blood flow signal passes an oscillatory cardiac signal componentand substantially rejects cardiac S1 and S2 signal peaks, thereby toseparate the varying blood flow signal.
 9. The electronic monitoringdevice claimed in claim 1 wherein the filter for said blood flow signalsmoothes the cardiac signal using a low-order polynomial filter with atime window approximately comparable in size to a period of oscillationof the varying blood flow signal.
 10. The electronic monitoring deviceclaimed in claim 9 wherein said electronic processor is operable tosubtract from the cardiac signal the result of the smoothing filter todeliver a residue signal primarily featuring the cardiac pulsesseparated from the cardiac signal.
 11. The electronic monitoring deviceclaimed in claim 10 wherein said electronic processor is operable tocount the cardiac pulses to deliver a heart rate signal.
 12. Theelectronic monitoring device claimed in claim 1 wherein the filter forsaid blood flow signal smooths the cardiac signal using anapproximately-4th order Savitzky-Golay polynomial filter with a windowsize approximately 200 milliseconds.
 13. The electronic monitoringdevice claimed in claim 1 wherein said electronic processor is operableto digitally low pass filter an input thereof with a cutoff lower than apower line frequency to isolate the cardiac signal.
 14. The electronicmonitoring device claimed in claim 1 further comprising at least twoanalog signal paths coupled to said electronic processor, each analogsignal path including an AC coupled amplifier feeding a low passanti-alias filter feeding a sampling analog-to-digital converter coupledto said electronic processor.
 15. The electronic monitoring deviceclaimed in claim 1 wherein the filter for said blood flow signalsmoothes the cardiac signal using a low-order polynomial filter toobtain the varying blood flow signal, and said electronic processorinstructions also define a substantially higher-order polynomial filterconcurrently operable to deliver a residue signal having cardiac pulses.16. The electronic monitoring device claimed in claim 1 wherein theelectronic processor is further operable to perform electronic peakdetection on the varying blood flow signal to identify a flow peakamplitude and a time location F1 of that peak.
 17. The electronicmonitoring device claimed in claim 16 wherein the electronic processoris further operable to separate the cardiac pulses and identify theirfirst peak location in time P1 for a same heart beat as applies to thetime location F1.
 18. The electronic monitoring device claimed in claim17 wherein the electronic processor is further operable to generate ahemodynamic parameter based on the time difference between time P1 andthe time F1.
 19. The electronic monitoring device claimed in claim 17wherein the electronic processor is further operable to generatebeat-by-beat values of the time difference between time P1 of the peakof an S1 cardiac pulse and the time F1 of the peak of the varying bloodflow signal.
 20. The electronic monitoring device claimed in claim 19wherein said electronic processor is further operable to derive arespiration-related signal from variations in the beat-by-beat values ofthe time difference.
 21. The electronic monitoring device claimed inclaim 1 wherein said electronic processor is further operable to performelectronic peak detection on the varying blood flow signal to identify aflow peak amplitude PAmp and generate a hemodynamic parameter as afunction of the flow peak amplitude PAmp.
 22. The electronic monitoringdevice claimed in claim 21 wherein said electronic processor is furtheroperable to derive a signal, related to at least one of respiration,intrapleural pressure, and intrathoracic pressure, from variations inbeat-by-beat values of the flow peak amplitude PAmp.
 23. The electronicmonitoring device claimed in claim 21 wherein said electronic processoris further operable to obtain the cardiac pulses and a heart rate HRtherefrom, and further to generate a hemodynamic parameter as a functionof the flow peak amplitude PAmp times the heart rate HR.
 24. Theelectronic monitoring device claimed in claim 1 further comprising amodem coupled to said electronic processor.
 25. The electronicmonitoring device claimed in claim 24 wherein said modem is selectedfrom the group consisting of: 1) cellular, 2) Wi-Fi, 3) wireline, 4)telephone modem.
 26. The electronic monitoring device claimed in claim 1wherein said electronic processor is also operable to substantiallyfilter out baseline-wander below about one-half Hertz from the bloodflow signal.
 27. An accelerometer sensor assembly having a broadsideportion and comprising an accelerometer sensor circuit having an axis ofacceleration sensitivity parallel to the broad side and orientable onthe chest to deliver an input signal representing a component ofacceleration approximately parallel to a head-to-feet body axis andincluding heart pulses and other variations mixed together; anelectronic circuit responsive to said accelerometer sensor circuit andoperable to execute a filter that delivers a varying blood flow signalfrom said input signal at least when that axis of accelerationsensitivity is approximately parallel to the head-to-feet body axis, thevarying blood flow signal substantially freed of heart pulses; and anoutput circuit operable to send information based on the varying bloodflow signal.
 28. The accelerometer sensor assembly claimed in claim 27further comprising a display coupled to said electronic circuit.
 29. Theaccelerometer sensor assembly claimed in claim 27 wherein saidaccelerometer sensor circuit has a second axis of accelerationsensitivity substantially perpendicular to the broad side and orientableon the chest to deliver a second input signal to said electronic circuitrepresenting a component of acceleration approximately parallel to adorsal-ventral direction.
 30. The accelerometer sensor assembly claimedin claim 27 wherein said accelerometer sensor circuit has another axisof acceleration sensitivity substantially parallel to the broad side andsubstantially perpendicular to the first-recited axis of accelerationsensitivity, and the electronic circuit coupled to receive another inputsignal from said accelerometer sensor circuit representing that otheraxis of acceleration sensitivity.
 31. The accelerometer sensor assemblyclaimed in claim 30 wherein said electronic circuit is operable tocombine signals from said accelerometer sensor circuit representing atleast both of those axes of acceleration sensitivity.
 32. Theaccelerometer sensor assembly claimed in claim 27 further comprising anelectrode for electrical skin contact and made physically part of theassembly.
 33. The accelerometer sensor assembly claimed in claim 27wherein said accelerometer sensor circuit includes at least two analogsignal conditioning paths.
 34. The accelerometer sensor assembly claimedin claim 27 wherein the varying blood flow signal has a spindle-shapedoscillatory waveform during each heartbeat and oscillating at afrequency higher than both a respiration frequency and a heart rate. 35.The accelerometer sensor assembly claimed in claim 27 wherein saidoutput circuit includes a modem coupled to said electronic circuit. 36.The accelerometer sensor assembly claimed in claim 27 wherein saidoutput circuit is selected from the group consisting of: 1) Bluetooth,2) Zigbee, 3) serial interface.
 37. The accelerometer sensor assemblyclaimed in claim 27 wherein said electronic circuit is operable toprovide a hemodynamic parameter, a heart rate, and a respirationparameter, all based on said accelerometer sensor circuit.
 38. Anelectronic monitoring process comprising: electronically bandpassingdigital signals representing living-body monitoring signals in a rangeselective for a cardiac signal; and executing a filter that delivers avarying blood flow signal from the cardiac signal.
 39. The electronicmonitoring process claimed in claim 38 further comprising separatingcardiac pulses from the cardiac signal and delivering an electronicrepresentation of the time interval between a cardiac pulse and an eventon the varying blood flow signal.
 40. The electronic monitoring processclaimed in claim 38 further comprising post-processing the varying bloodflow signal as from a 2nd order system model.
 41. The electronicmonitoring process claimed in claim 38 further comprisingpost-processing the varying blood flow signal and the cardiac signalinto electronically-represented displays of physiological function. 42.A hemodynamic monitoring device comprising: an electronic processorhaving at least one signal input for body monitoring; and a memoryholding instructions for said electronic processor coupled to saidelectronic processor so that said electronic processor is operable toobtain a cardiac pulse signal having a varying amplitude, and saidelectronic processor further operable to post-process the cardiac pulsesignal to provide a time-varying output representing an estimation forat least one hemodynamic parameter based on the amplitude and selectedfrom the group consisting of stroke volume SV and cardiac output CO. 43.The hemodynamic monitoring device claimed in claim 42 wherein saidelectronic processor is also operable to separate a respiration signalalso based on the amplitude from the cardiac pulse signal.
 44. Thehemodynamic monitoring device claimed in claim 42 further comprising adisplay coupled to said electronic processor to display the at least onehemodynamic parameter.